Devices and methods for tracking blood flow and determining parameters of blood flow

ABSTRACT

Provided herein is a method for use in medical applications that permits (1) affordable three-dimensional imaging of blood flow using a low-profile easily-attached transducer pad, (2) real-time blood-flow vector velocity, and (3) long-term unattended Doppler-ultrasound monitoring in spite of motion of the patient or pad. The pad and associated processor collects and Doppler processes ultrasound blood velocity data in a three dimensional region through the use of a planar phased array of piezoelectric elements. The invention locks onto and tracks the points in three-dimensional space that produce the locally maximum blood velocity signals. The integrated coordinates of points acquired by the accurate tracking process is used to form a three-dimensional map of blood vessels and provide a display that can be used to select multiple points of interest for expanded data collection and for long term continuous and unattended blood flow monitoring. The three dimensional map allows for the calculation of vector velocity from measured radial Doppler. 
     A thinned array (greater than half-wavelength element spacing of the transducer array) is used to make a device of the present invention inexpensive and allow the pad to have a low profile (fewer connecting cables for a given spatial resolution). The full aperture is used for transmit and receive so that there is no loss of sensitivity (signal-to-noise ratio) or dynamic range. Utilizing more elements (extending the physical array) without increasing the number of active elements increases the angular field of view. A further increase is obtained by utilizing a convex non-planar surface.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.10/831,547 filed Apr. 23, 2004, which is a continuation in part of U.S.application Ser. No. 10/327,265 filed Dec. 20, 2002, which is acontinuation-in-part of U.S. application Ser. No. 09/926,666, now U.S.Pat. No. 6,682,483, which is a National Phase Application ofInternational Application No. PCT/US00/14691 filed May 26, 2000 and isbased upon U.S. Provisional Application Nos. 60/136,364, filed 28 May1999; 60/138,793, filed 14 Jun. 1999; and 60/152,886, filed Sep. 8,1999. The above-referenced U.S. application Ser. No. 10/327,265 claimspriority to U.S. provisional application Ser. No. 60/343,061.

FIELD OF THE INVENTION

The present invention involves an ultrasound Doppler method that permitsnon-invasive diagnosis and non-invasive unattended, continuousmonitoring of vascular blood flow for medical applications. Theinvention further relates to a method and a system for the digital beamforming at the transmit and/or receive side of a phased arraytransducer.

BACKGROUND OF THE INVENTION

Blood velocity monitoring is not currently practical for intensive careunit (ICU) or surgical applications. For non-invasive brain bloodvelocity monitoring, for example, a transcranial Doppler (TCD) probemust be mounted in a ball joint that is attached to the head by ahelmet. The probe must be carefully aimed and fastened in place by anexperienced person who knows how to locate the middle cerebral artery.Slight movements cause the probe to lose the blood velocity signal.Moreover, conventional Doppler ultrasound probes used in these devicesscan (either mechanically or by using an acoustic phased array) in onlyone angle (which we will call azimuth), and will map only a single sliceof the object being imaged.

Efforts have been made to modify such devices to provide real-time threedimensional (3-D) imaging. However, in order for a two dimensional (2-D)device to provide such imaging normally requires thousands of elements,and must form many thousands of pencil beams every 1/30 second. Sensorcost grows with the number of elements in the array and the number ofprocessing channels. Thus, such devices are cost prohibitive, as well asimpractical.

Moreover, no automated procedure exists in current practice forprecisely locating the optimum point at which to measure the Dopplersignal. Conventional ultrasound Doppler-imaging devices can only measureradial velocity in blood vessels, and not the vector velocity ormagnitude of the velocity of the blood.

Accordingly, what is needed is a new and useful Doppler ultrasounddevice and method that can automatically locate the optimum point atwhich to measure the Doppler signal, and thus provide medical providerswith parameters such as vector velocity, the volume of blood passingthrough the blood vessel and the Doppler spectral distribution of theblood flow.

What is also needed is a new and useful Doppler ultrasound device andmethod that does not require it be placed on a patient with precision,and will enable a patient wearing the device to move freely.

The citation of any reference herein should not be construed as anadmission that such reference is available as “Prior Art” to the instantapplication.

SUMMARY OF THE INVENTION

There is provided, in accordance with the present invention, a new,useful, and—unobvious method of determining parameters of blood flow,such as vector velocity, blood flow volume, and Doppler spectraldistribution, using sonic energy (ultrasound) and a novel thinned array.Also provided is a novel method of tracking blood flow and generating athree dimensional image of a blood vessel of interest that has muchgreater resolution than images produced using heretofore knownultrasound devices and methods.

Broadly, the present invention extends to a method for determining aparameter of blood flow in a blood vessel of interest, comprising thesteps of:

-   -   a) providing an array of sonic transducer elements, wherein the        element spacing in the array is greater than, equal or less than        a half wavelength of the sonic energy produced by the elements,        wherein at least one element transmits sonic energy, and a        portion of the elements receive sonic energy;    -   b) directing sonic energy produced by the at least one element        of the array into a volume of the subject's body having the        blood vessel of interest,    -   c) receiving echoes of the sonic energy from the volume of the        subject's body having the blood vessel of interest:    -   d) reporting the echoes to a processor programmed to:        -   i) Doppler process the echoes to determine radial velocity            of the blood flowing in the blood vessel of interest;        -   ii) calculate a three dimensional position of blood flow in            the vessel of interest; and        -   iii) calculate the parameter of blood flow in the blood            vessel at the three dimensional position calculated in step            (ii); and    -   (e) displaying the parameter on a display monitor that is        electrically connected to the processor.

Moreover, a method of the present invention permits an operatorexamining a subject to obtain information on blood flow in a particularregion of the blood vessel of interest.

As used herein, the phrases ˜element spacing” and “distance between theelements” can be used interchangeably and refer to the distance betweenthe center of elements of an array.

Various methods can be used to determine the three, dimensional positionof blood flow. In a particular embodiment, the method comprises thesteps of having the processor programmed to:

-   -   i) determine a sum beam, an azimuth difference beam and an        elevation difference beam from the echoes received from the        blood vessel of interest;    -   ii) modulate the directions of the transmitted and received        sonic energy based upon the sum, azimuth difference and        elevation difference beams in order to lock on to the highest        Doppler energy calculated from echoes from the flow of blood in        the blood vessel of interest, and    -   iii) calculate the three dimensional position of the highest        Doppler energy from the blood flow in the vessel of interest.

Optionally, the processor can also be programmed to determine at leastone additional beam having an angle between the azimuth difference beamand the elevation difference beam prior to modulating the directions ofthe transmitted and received sonic energy, wherein the at least oneadditional beam is used to modulate the directions of the transmittedand received sonic energy. Naturally, the angle of the at least oneadditional beam can vary. In a particular embodiment, the at least oneadditional beam is at an angle that is orthogonal to the blood vessel ofinterest.

Moreover, the present invention extends to a method as described above,wherein steps (b) through (e) are periodically repeated so that thethree dimensional position of blood flow in the vessel of interest istracked, and the parameter of blood flow is periodically calculated anddisplayed on the display monitor. In a particular embodiment, the periodof time between repeating steps (b) through (e) is sufficiently short sothat the parameter being measured remains constant, e.g., 20milliseconds.

The present invention further extends to a method for determining aparameter of blood flow in a particular region of a blood vessel ofinterest, comprising the steps of:

-   -   a) providing an array of sonic transducer elements, wherein the        element spacing in the array is greater than, equal or less than        a half wavelength of the sonic energy produced by the elements,        wherein at least one element transmits sonic energy, and a        portion of the elements receive sonic energy;    -   b) directing sonic energy produced by the at least one element        of the array into a volume of the subject's body having the        particular region of the blood vessel of interest,    -   c) receiving echoes of the sonic energy from the volume of the        subject's body having the particular region of the blood vessel        of interest;    -   d) reporting the echoes to a processor programmed to        -   i) Doppler process the echoes to determine radial velocity            Of the blood flowing in the particular region of the blood            vessel of interest;        -   ii) calculate a three dimensional position of blood flow in            the particular region of the blood vessel of interest; and        -   iii) calculate the parameter of blood flow in the particular            region of the blood vessel of interest at the three            dimensional position calculated in step (ii); and    -   (e) displaying the parameter on a display monitor that is        electrically connected to the processor.

A particular method of calculating the three dimensional position ofblow flow in such a method of the present invention comprises having theprocessor programmed to:

-   -   i) determine a sum beam, an azimuth difference beam and an        elevation difference beam from the echoes received from the        particular region of the blood vessel of interest;    -   ii) modulate the directions of the transmitted and received        sonic energy based upon the sum, azimuth difference and        elevation difference beams in order to lock on to the highest        Doppler energy calculated from echoes received from the flow of        blood in the particular region of the blood vessel of interest,        and    -   iii) calculate the three dimensional position of the highest        Doppler energy from the blood flow in the particular region of        the blood vessel of interest.

As explained above, at least one additional beam can also be determinedand used to calculate the three dimensional position.

Furthermore, the present invention extends to a method for determining aparameter of blood flow in a blood vessel of interest, comprising thesteps of:

-   -   (a) providing an array of sonic transducer elements, wherein the        element spacing in the array is greater than, equal or less than        a half wavelength of the sonic energy produced by the elements,        wherein at least one element transmits sonic energy, and a        portion of elements receive sonic energy;    -   (b) directing sonic energy produced by the at least one element        of the array into a volume of the subject's body having the        blood vessel of interest,    -   (c) receiving echoes of the sonic energy from the volume of the        subject's body having the blood vessel of interest;    -   d) reporting the echoes to a processor electrically connected to        the elements of the array, wherein the processor is programmed        to:        -   i) Doppler process the echoes to determine radial velocity            of the blood flowing in the blood vessel of interest;        -   ii) determine a sum beam, an azimuth difference beam and an            elevation difference beam from the echoes received from the            blood vessel of interest;        -   iii) modulate the directions of the transmitted and received            sonic energy based upon the sum, azimuth difference and            elevation difference beams in order to lock on to the            highest Doppler energy calculated from echoes from the flow            of blood in the blood vessel of interest,        -   iv) calculate the three dimensional position of the highest            Doppler energy from the blood flow in the vessel of            interest; and        -   v) calculate the parameter of blood flow in the blood vessel            at the three dimensional position calculated in step (iv);            and    -   (e) displaying the parameter on a display monitor that is        electrically connected to the processor.

As explained above, an operator performing a method of the presentinvention can obtain blood flow parameters from a blood vessel ofinterest, and even from a particular region of a blood vessel ofinterest.

Moreover, the present invention extends to a method for determining aparameter of blood flow in a particular region of a blood vessel ofinterest, comprising the steps of:

-   -   a) providing an array of sonic transducer elements, wherein the        element spacing in the array is greater than, equal or less than        a half wavelength of the sonic energy produced by the elements,        wherein at least one element transmits sonic energy, and a        portion of the elements receive sonic energy;    -   b) directing sonic energy produced by the at least one element        of the array into a volume of the subject's body having the        particular region of the blood vessel of interest,    -   c) receiving echoes of the sonic energy from the volume of the        subject's body having the particular region of blood vessel of        interest;    -   d) reporting the echoes to a processor electrically connected to        the elements of the array, wherein the processor is programmed        to        -   i) Doppler process the echoes to determine radial velocity            of the blood flowing in the particular region of the blood            vessel of interest;        -   ii) determine a sum beam, an azimuth difference beam and an            elevation difference beam from the echoes received from the            particular region of the blood vessel of interest;        -   iii) modulate the directions of the transmitted and received            sonic energy based upon the sum, azimuth difference and            elevation difference beams in order to lock on to the            highest Doppler energy calculated from echoes from the flow            of blood in the particular region of the blood vessel of            interest,        -   iv) calculate the three dimensional position of the highest            Doppler energy from the blood flow in the particular region            of the blood vessel of interest; and        -   v) calculate the parameter of blood flow in the particular            region of the blood vessel at the three dimensional position            calculated in step (iv); and    -   (e) displaying the parameter on a display monitor that is        electrically connected to the processor.

In another embodiment, the present invention extends to a device fordetermining a parameter of blood flow in a blood vessel of interest,comprising:

-   -   (a) an array of sonic transducer elements, wherein the element        spacing in the array is greater than, equal or less than a half        wavelength of the sonic energy produced by the elements, and at        least one element transmits sonic energy and a portion of the        elements receive sonic energy;    -   (b) a processor electrically connected to the array so that        echoes received from a volume of the subject's body having the        blood vessel of interest due to directing sonic energy produced        by the at least one element of the array into the subject's body        is reported to the processor, wherein the processor is        programmed to:        -   i) Doppler process the echoes to determine radial velocity            of the blood flowing in the blood vessel of interest;        -   ii) calculate a three dimensional position of blood flow in            the blood vessel of interest; and        -   iii) calculate the parameter of blood flow in the blood            vessel of interest at the three dimensional position            calculated in step (ii); and    -   (c) a display monitor that is electrically connected to the        processor which displays the parameter of blood flow calculated        by the processor.

A parameter of blood that can be determined with a device of the presentinvention includes blood flow volume, vector velocity, Doppler spectraldistribution, etc. The parameter being measured can be an instantaneousvalue, or an average value determined over a heart cycle.

Moreover, the present invention extends to a device as described above,wherein the processor is programmed to:

-   -   i) determine a sum beam, an azimuth difference beam and an        elevation difference beam from the echoes received from the        blood vessel of interest after Doppler processing the echoes;    -   ii) modulate the directions of the transmitted and received        sonic energy based upon the sum, azimuth difference and        elevation difference beams in order to lock on to the highest        Doppler energy calculated from echoes from the flow of blood in        the blood vessel of interest,    -   iii) calculate the three dimensional position of the highest        Doppler energy from the blood flow in the vessel of interest;        and    -   iv) calculate the parameter of blood flow in the blood vessel of        interest at the three dimensional position calculated in (iii).

Optionally, a processor of a device of the present invention can befurther programmed to determine at least one additional beam having anangle between the azimuth difference beam and the elevation differencebeam prior to modulating the directions of the transmitted and receivedsonic energy, wherein the at least one additional beam is used tomodulate the directions of the transmitted and received sonic energy. Ina particular embodiment, the at least one additional beam is at an anglethat is orthogonal to the blood vessel of interest.

Moreover, in another embodiment of a device of the present invention,the distance between the elements of the array is greater than ½ thewavelength of the sonic energy generated by the at least one element.

Furthermore, the present invention extends to a device for determining aparameter of blood flow in a blood vessel of interest, comprising:

-   -   (a) an array of sonic transducer elements, wherein the element        spacing in the array is greater than, equal or less than a half        wavelength of the sonic energy produced by the elements, and at        least one element transmits sonic energy, and portion of the        elements receive sonic energy;    -   b) processor electrically connected to the array so that echoes        received from a volume of the subject's body having the blood        vessel of interest due to directing sonic energy produced by the        at least one element of the array into the subject's body is        reported to the processor, wherein the processor is programmed        to:        -   i) Doppler process the echoes to determine radial velocity            of the blood flowing in the blood vessel of interest;        -   ii) calculate a three dimensional position of blood flow in            the blood vessel of interest; and        -   iii) calculate the parameter of blood flow in the blood            vessel of interest at the three dimensional position            calculated in step (ii)    -   (c) a display monitor that is electrically connected to the        processor which displays the parameter of blood flow calculated        by the processor.

Particular parameters of blood flow that can be determined with a deviceof the present invention include, but certainly are not limited to bloodflow volume, vector velocity, and Doppler spectral distribution. Theparameter being measured can be an instantaneous value, or an averagevalue determined over a heart cycle.

In addition, a processor of a device of the present invention can befurther programmed to determine at least one additional beam having anangle between the azimuth difference beam and the elevation differencebeam prior to modulating the directions of the transmitted and receivedsonic energy, wherein the at least one additional beam is used tomodulate the directions of the transmitted and received sonic energy. Ina particular embodiment, the at least one additional beam is at an anglethat is orthogonal to the blood vessel of interest.

Moreover, the present invention extends to a method for generating athree dimensional image using sonic energy of a blood vessel of interestin a subject, the method comprising the steps of:

-   -   (a) providing an array of sonic transducer elements, wherein the        element spacing in the array is greater than, equal or less than        a half wavelength of the sonic energy produced by the elements,        wherein at least one element transmits sonic energy, and a        portion of the elements receive sonic energy;    -   (b) directing sonic energy produced by the at least one element        of the array into a volume of the subject's body having the        blood vessel of interest,    -   (c) receiving echoes of the sonic energy from the volume of the        subject's body having the blood vessel of interest;    -   (d) reporting the echoes to a processor programmed to        -   i) Doppler process the echoes to determine radial velocity            of the blood flowing in the blood vessel of interest;        -   ii) calculate a three dimensional position of blood flow in            the blood vessel of interest;        -   iii) repeat steps (i) through (ii) to generate a plurality            of calculated three dimensional positions; and        -   vi) generate a three dimensional image of the blood vessel            of interest from the plurality of calculated three            dimensional positions; and    -   (e) displaying the three dimensional image on a display monitor        that is electrically connected to the processor.

Furthermore, the present invention permits an operator utilizing amethod of the present invention to generate a three dimensional image ofnot only a blood vessel in the body, but even a particular region of ablood vessel in the body.

Numerous means available for calculating the three dimensional positionof a blood vessel and even a particular portion of a blood vessel areencompassed by the present invention. A particular means compriseshaving the programmed processor:

-   -   i) determine a sum beam, an azimuth difference beam and an        elevation difference beam from the echoes received from the        blood vessel of interest after Doppler processing the echoes;    -   ii) modulate the directions of the transmitted and received        sonic energy based upon the sum, azimuth difference and        elevation difference beams in order to lock on to the highest        Doppler energy calculated from echoes from the flow of blood in        the blood vessel of interest, and    -   iii) calculate the three dimensional position of the highest        Doppler energy from the blood flow in the vessel of interest,        and    -   iv) repeat steps (i) through (iii) to generate a plurality of        calculated three dimensional positions.

Optionally, a processor of a method of the present invention can also beprogrammed to determine at least one additional beam having an anglebetween the azimuth difference beam and the elevation difference beamprior to modulating the directions of the transmitted and received sonicenergy, and the at least one additional beam is also used to modulatethe directions of the transmitted and received sonic energy, andcalculate the three dimensional position of the highest Doppler energy.In a particular embodiment, the at least one additional beam is at anangle that is orthogonal to the blood vessel of interest.

The present invention also extends to a method for generating a threedimensional image of a blood vessel of interest in a subject using sonicenergy, the method comprising the steps of:

-   -   (a) providing an array of sonic transducer elements, wherein the        element spacing in the array is greater than, equal or less than        a half wavelength of the sonic energy produced by the elements,        wherein at least one element transmits sonic energy, and a        portion of the elements receive sonic energy;    -   (b) directing sonic energy produced by the at least one element        of the array into a volume of the subject's body having the        blood vessel of interest,    -   (c) receiving echoes of the sonic energy from the volume of the        subject's body having the blood vessel of interest;    -   (d) reporting the echoes to a processor programmed to        -   i) Doppler process the echoes to determine radial velocity            of the blood flowing in the blood vessel of interest;        -   ii) determine a sum beam, an azimuth difference beam and an            elevation difference beam from the echoes received from a            portion of the blood vessel of interest;        -   ii) modulate the directions of the transmitted and received            sonic energy based upon the sum, azimuth difference and            elevation difference beams in order to lock on to the            highest Doppler energy calculated from echoes from the flow            of blood in the blood vessel of interest,        -   iv) calculate the three dimensional position of the highest            Doppler energy from the blood flow in the vessel of            interest; and        -   v) repeat steps (i) through (iv) to generate a plurality of            calculated three dimensional positions;        -   vi) generate a three dimensional image of the blood vessel            of interest from the plurality of calculated three            dimensional positions; and    -   (e) displaying the three dimensional image on a display monitor        that is electrically connected to the processor.

Optionally, the three dimensional image can be of a particular region ofa blood vessel of interest. Moreover, a processor of a method describedherein can also determine at least one additional beam having an anglebetween the azimuth difference beam and the elevation difference beamprior to modulating the directions of the transmitted and received sonicenergy, and the at least one additional beam is also used to modulatethe directions of the transmitted and received sonic energy, andcalculate the three dimensional position of the highest Doppler energy.Angles for use with the at least one additional beam are describedabove.

Moreover, in another embodiment of the present invention, the distancebetween the elements of the array is greater than ½ the wavelength ofthe sonic energy generated by the at least one element.

Furthermore, the present invention extends to a device generating athree dimensional image of a blood vessel of interest in a subject usingsonic energy, comprising:

-   -   (a) an array of sonic transducer elements, wherein the element        spacing in the array is greater than, equal or less than a half        wavelength of the sonic energy produced by the elements, and at        least one element transmits sonic energy, and a portion of the        elements receive sonic energy;    -   (b) a processor electrically connected to the array so that        echoes received from a volume of the subject's body having the        blood vessel of interest due to directing sonic energy produced        by the at least one element of the array into the subject's body        is reported to the processor, wherein the processor is        programmed to:        -   i) Doppler process the echoes to determine radial velocity            of the blood flowing in the blood vessel of interest;        -   ii) calculate a three dimensional position of blood flow in            the blood vessel of interest;        -   iii) repeat steps (i) through (ii) to generate a plurality            of calculated three dimensional positions;        -   v) generate a three dimensional image from the plurality of            calculated three dimensional positions, and    -   (c) a display monitor that is electrically connected to the        processor which displays the three dimensional image.

As explained above, a device of the present invention permits anoperator to generate and display three dimensional images of a bloodvessel of interest, and even of a particular region of a blood vesselthat the operator wants to investigate closely. Moreover, in aparticular embodiment, a processor of a device of the present inventioncan be programmed to calculate the three dimensional position of a bloodvessel by:

-   -   i) determining a sum beam, an azimuth difference beam and an        elevation difference beam from the echoes received from the        blood vessel of interest after Doppler processing the echoes;    -   ii) modulating the directions of the transmitted and received        sonic energy based upon the sum, azimuth difference and        elevation difference beams in order to lock on to the highest        Doppler energy calculated from echoes from the flow of blood in        the blood vessel of interest,    -   iii) calculating. the three dimensional position of the highest        Doppler energy from the blood flow in the vessel of interest;        and    -   iv) repeat steps (I) through (iii) in order to generate a        plurality of calculated three dimensional positions used to        generate the three dimensional image.

Optionally, the processor can be programmed to further determine atleast one additional beam having an angle between the azimuth differencebeam and the elevation difference beam prior to modulating thedirections of the transmitted and received sonic energy, wherein the atleast one additional beam is used to modulate the directions of thetransmitted and received sonic energy. The angle between the azimuthdifference beam and the elevation difference beam of the additional beamcan vary. In a particular embodiment, the at least one additional beamis at an angle that is orthogonal to the blood vessel of interest.

Furthermore, the present invention extends to a thinned array for use inan ultrasound device, comprising a plurality of sonic transducerelements, wherein the element spacing in the array is greater than ahalf wavelength of the sonic energy produced by the elements, and theelements are positioned and sized within the array, and sonic energy iselectronically steered by the elements so that any grating lobesproduced by the sonic energy are suppressed. In a particular embodiment,the elements positioned and sized so that they are flush against eachother.

Hence, the current invention performs blood velocity monitoring bycollecting Doppler data in three dimensions; azimuth, elevation, andrange (depth); so that the point (in three dimensional space) at whichthe velocity is to be monitored can be acquired and tracked when thepatient or the sensor moves. The invention also produces a threedimensional map of the blood flow and converts measured radial velocityto true vector velocity.

Moreover, in this invention, once the desired signal is found, it willbe precisely located and continually tracked with accuracy far betterthan the resolution. A heretofore unknown method to achievesub-resolution tracking and mapping involves a novel and unobviousextension of a procedure called “monopulse”. Monopulse tracking has beenused in military applications for precisely locating and tracking apoint target with electromagnetic radiation. However, it has never beenutilized in connection with sonic waves to determine the velocity ofmoving fluids in vivo.

This invention provides: (1) affordable three-dimensiortal imaging ofblood flow using a low-profile easily-attached transducer pad, (2)real-time vector velocity, and (3) long-term unattendedDoppler-ultrasound monitoring in spite of motion of the patient or pad.None of these three features are possible with current ultrasoundequipment or technology.

The pad and associated processor collects and Doppler processesultrasound blood velocity data in a three-dimensional region through theuse of a two-dimensional phased array of piezoelectric elements on aplanar, cylindrical, or spherical surface:

Through use of unique beamforming and tracking techniques, the inventionlocks onto and tracks the points in three-dimensional space that producethe locally maximum blood velocity signals. The integrated coordinatesof points acquired by the accurate tracking process is used to form athree-dimensional map of blood vessels and provide a display that can beused to select multiple points of interest for expanded data collectionand for long term continuous and unattended blood flow monitoring. Thethree dimensional map allows for the calculation of vector velocity frommeasured radial Doppler.

In a particular embodiment, a thinned array (greater than halfwavelength element spacing of the transducer array), is used to make adevice of the present invention inexpensive and allow the pad to have alow profile (fewer connecting cables for a given spatial resolution).The array is thinned without reducing the receiver area by limiting theangular field of view. The special 2-0 phased array used in thisinvention makes blood velocity monitoring inexpensive and practical by(1) forming the beams needed for tracking and for re-acquiring the bloodvelocity signal and by (2) allowing for an element placement that issignificantly coarser than normal half-wavelength element spacing. Thelimited range of angles that the array must search allows for much lessthan the normal half wavelength spacing without reducing the totalreceiver area.

Grating lobes due to array thinning can be reduced by using widebandwidth and time delay steering. The array, or at least one element ofthe array, is used to sequentially insonate the beam positions. Once theregion of interest has been imaged and coarsely mapped, the array isfocused at a particular location on a particular blood vessel formeasurement and tracking. Selection of the point or points to bemeasured and tracked can be based on information obtained via mappingand may be user guided or fully automatic. Selection can be based, forexample, on peak response within a range of Doppler frequencies at ornear an approximate location.

In the tracking mode a few receiver beams are formed at a time: sum,azimuth difference, elevation difference, and perhaps, additionaldifference beams, at angles other than azimuth (0 degrees) and elevation(=90 degrees). Monopulse is applied at angles other than 0 and 90degrees (for example 0, 45, 90, and 135 degrees) in order to locate avessel in a direction perpendicular to the vessel. When the desired(i.e. peak) blood velocity signal is not in the output, this isinstantly recognized (e.g., a monopulse ratio, formed after Dopplerfiltering, becomes nonzero) and the array is used to track (slowmovement) or re-acquire (fast movement) the desired signal.Re-acquisition is achieved by returning to step one to form andDoppler-process a plurality of beams in order to select the beam (andthe time delay or “range gate”) with the most high-Doppler (high bloodvelocity) energy. This is followed by post-Doppler monopulse tracking tolock a beam and range gate on to the exact location of the peak velocitysignal. In applications such as transcranial Doppler, where angularresolution based on wavelength and aperture size is inadequate, finemapping is achieved, for example, by post-Doppler monopulse trackingeach range cell of each vessel, and recording the coordinates andmonopulse-pair angle describing the location and orientation of themonopulse null. With a three-dimensional map available, true vectorvelocity can be computed. For accurate vector flow measurement, themonopulse difference is computed in a direction orthogonal to the vesselby digitally rotating until a line in the azimuth-elevation or C-scandisplay is parallel to the vessel being monitored. The aperture is moreeasily rotated in software (as opposed to physically rotating thetransducer array) if the aperture is approximately circular (oreliptical) rather than square (or rectangular). Also, lower sidelobesresult by removing elements from the four corners of a square orrectangular array in order to make the array an octagon.

In this invention, as long as (1) a blood vessel or (2) a flow region ofa given velocity can be resolved by finding a 3-D resolution cellthrough which only a single vessel passes, that vessel or flow componentcan then be very accurately located within the cell. Monopulse is merelyan example of one way to attain such sub-resolution accuracy (SRA).Other methods involve “super-resolution” or “parametric” techniques usedin “modern spectral estimation”, including the MUSIC algorithm andautoregressive modeling, for example. SRA allows an extremely accuratemap of 3-D flow.

Furthermore, the present invention utilizes post-Doppler, sub-resolutiontracking and mapping; it does Doppler processing first and uses onlyhigh Doppler-frequency data.

This results in extended targets since the active vessels approximate“lines” as opposed to “points”. In three-dimensional space, thesevessels are resolved, one from another. At a particular range, themonopulse angle axis can be rotated (in the azimuth-elevation plane) sothat the “line” becomes a “point” in the monopulse angle direction. Thatpoint can then be located by using super-resolution techniques or byusing a simple technique such as monopulse. By making many suchmeasurements an accurate 3-D map of the blood vessels results.

Methods for extending the angular field of view of the thinned array(that is limited by grating lobes) include (1) using multiple panels oftransducers with multiplexed processing channels. (2) convex V-shapedtransducer panels, (3) cylindrical shaped transducer panel, (4)spherical shaped transducer panel, and (5) negative ultrasound lens. Ifneeded, moving the probe and correlating the sub-images can create a mapof an even larger region.

Active digital beamforming can also be utilized, but the implementationdepends on a choice to be made between wideband and narrowbandimplementations. If emphasis is on high resolution mapping of the bloodvessels, then a wide bandwidth (e.g., 50% of the nominal frequency) isused for fine range resolution. If emphasis is on Doppler spectralanalysis, measurement, and monitoring, the map is only a tool. In thiscase, a narrowband, low cost, low range-resolution, high sensitivityimplementation might be preferred. A wideband implementation wouldbenefit in performance (higher resolution, wider field of view, andreduced grating lobes) using time-delay steering while a narrowbandimplementation would benefit in cost using phase-shift steering. Theinvention can thus be described in terms of two preferredimplementations.

In a wideband implementation, time delay steering can be implementeddigitally for both transmit and receive by over-sampling and digitallydelaying in discrete sample intervals. In a narrowband implementation,(1) phase steering can be implemented digitally (digital beamforming)for both transmit and receive, and (2) bandpass sampling (sampling at arate lower than the signal frequency) can be employed with digitaldown-conversion and filtering.

Accordingly, it is an object of the present invention to locate thepoint in three dimensional space having the greatest high-Dopplerenergy, and determining coordinates for that point. With thatinformation, and the radial velocity of the blood flowing through theblood vessel at that point, a variety of blood flow parameters can becalculated at that point, including, but not limited to vector velocityof blood flow, volume of blood flow, or Doppler spectral distribution.The parameter being measured can be an instantaneous value, or anaverage value determined over a heart cycle.

It is also an object of the present invention to continuously track andmap in vivo the point in three dimensional space having the greatestDoppler-energy, and using the coordinates to generate a threedimensional image of a blood vessel and blood flow therein that possessa much greater resolution than images generated using heretofore knownDoppler ultrasound methods and devices.

It is yet another object of the present invention to provide a thinnedarray which does not utilize the number of element transducers as arerequired with heretofore known Doppler ultrasound devices. As a result,the decreased number of elements in the array decreases size of thearray utilized and provides a patient being analyzed with mobility thatwould not be available if using conventional ultrasound devices toobtain blood flow parameters such as vector velocity, blood flow volume,and Doppler spectral distribution. The parameter being measured can bean instantaneous value, or an average value determined over a heartcycle.

These and other aspects of the present invention will be betterappreciated by reference to the following drawings and DetailedDescription.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates the Blood Flow Mapping Monitor in use with aTranscranial Doppler Probe, as an example.

FIG. 2 shows a 64-element bistatic ultrasound transducer array example,where, with D=2d, the same elements are reconfigured differently fortransmit and receive during the acquisition phase of operation. FIG. 2(a) shows the Receive Configuration, where all 64 elements receive atonce. FIG. 2 (b) shows the Transmit Configuration, where, duringacquisition. the 16 sub-apertures transmit one at a time.

FIG. 3 is an example overall block diagram of a blood flow mappingmonitor embodiment.

FIG. 4 illustrates ultrasound beam coverage for the TCD array example ofFIG. 2. The left illustration shows 25 digitally beam-formed beams, asan example. On the right, is shown, for that example, the manner inwhich the transmit beam encompasses 21 receive beams in the acquisitionmode.

FIG. 5 shows one-dimensional patterns for a bistatic transducer arraywith D=2 d as in FIG. 2. FIG. 5 a (top) shows the transmit elementpattern. FIG. 5 b shows the receive Element Pattern and Array Patternwith the receiver beam steered to broadside (x=0). The Array Pattern hasGrating Lobes (Receiver Ambiguities). FIG. 5 c shows the resultanttwo-way beam pattern (product of all three patterns above). The GratingLobes are suppressed.

FIG. 6 is the same as FIG. 5, with the receive array beam steered tox=0.2.

FIG. 7 shows the Two-way pattern of a receiver beam steered to the halfpower point (x=0.2). This is FIG. 6 c plotted in dB.

FIGS. 8A-C show a one-dimensional representation of the example of FIG.4. FIG. 8A shows the product of transmit and receive Element Patterns.FIG. 8B plots a set of five receive beams showing Grating Lobes of theThinned Array. FIG. 8C plots the resultant two-way beams with GratingLobes suppressed.

FIG. 9 is a block diagram of one possible embodiment of theTransmit-Receive Electronics for a Bistatic Ultrasound Imaging Sensorand Blood Monitoring Monitor.

FIG. 10 shows the receiver channel signal spectrum illustratingfunctions performed by the FPGA of FIG. 9 on each of the 64 receivedsignals for a narowband case.

FIG. 11 shows the geometry involved in using azimuth monopulse to moreaccurately determine the cross-range location of a vessel. The rangeresolution is better than the cross-range resolution and the measuredradial velocity field or color flow map has been utilized to rotate andorient the azimuth and elevation axes so that the center of the vesselis vertical, at approximately zero azimuth. The black circular cylinderrepresents the location of all points within the spatial resolution cellthat have a particular velocity.

FIG. 12 shows the geometry involved in using Doppler ultrasound todetermine the diameter of a vessel or the velocity field within thevessel. While the initial 3-D orientation of the vessel is general, ameasured 3-D radial velocity field or 3-D color flow map has beenutilized to rotate and orient the azimuth and elevation axes so that thecenter of the vessel is vertical, at approximately zero azimuth. Inother words, the coordinate system has been rotated about the depth-axisso that the centerline of the vessel is in the depth-elevation plane.This can be accomplished either by a change of coordinates in softwareor by physically rotating the ultrasound probe. The black circularcylinder represents the location of all points within the illustratedbox that have a particular velocity. The diameter of the cylinder isthen measured as the azimuth extent of a high-resolution depth-azimuthor B-scan image at the Doppler frequency under examination.

FIG. 13 illustrates the Blood Flow Mapping Monitor in use with aTranscranial Doppler Probe, as an example.

FIG. 14 shows a 52-element ultrasound transducer array example, based onan 8 by 8 rectangular array of elements with 3 elements removed fromeach corner to make the array octagonal instead of rectangular orsquare. For this example, the elements are square (d₁=d₂=d) and L/d=8.

FIG. 15 shows a typical pattern of electronically scanned beams producedby the array in FIG. 14. The beam width is nominally, given by thesignal wavelength divided by the size, L, of the array. The angularfield of view (F.O.V.) is limited by the maximum angle to which thearray can be steered without producing grating lobes that are notsufficiently attenuated by the pattern of the individual d×d element.

FIGS. 16A-B show one-dimensional patterns for an eight-elementmonostatic linear transducer array corresponding to a column or a row inFIG. 16. FIG. 16A (top) shows the Element Pattern and Array Pattern withthe beam steered to broadside (x=0). The Array Pattern has Grating Lobes(Receiver Ambiguities). FIG. 16B shows the resultant beam pattern. TheGrating Lobes are suppressed.

FIGS. 17A-B are the same as FIGS. 16A-B, with the array beam steered toan angle at which a grating lobe exceeds the highest sidelobe. Thethinned array of FIG. 16 should not be steered beyond ±arcsin (λ/5d)(+4.7° for the example used) if grating lobes are to be suppressed.

FIG. 18 shows the pattern of a beam steered to the point where thegrating lobe problem appears. This is FIG. 17 b plotted in dB.

FIG. 19A shows a dual 52-active-element ultrasound transducer arrayexample (similar to that in FIG. 14) with a total of 116 elements, 52 ofwhich are used at a time. FIG. 19 B shows that the two sub-arrays are intwo different planes, tilted to reduce the overlap between beams fromthe two sub-arrays and maximize the azimuth angular field of view.

FIG. 20 shows a 52-active-element ultrasound transducer array example(similar to that in FIG. 14) with a total of 84 elements (52 of whichare used at a time) and with a slightly convex cylindrical shape. Theindicated L₁×L₂ sub-aperture would be activated for the formation ofbeams pointed to one side.

FIG. 21 is an example overall block diagram of a blood flow mappingmonitor embodiment.

FIG. 22 is a block diagram of one possible embodiment of the analogTransmit-Receive Electronics for an Ultrasound Imaging Sensor and BloodMonitor.

FIG. 23 shows the geometry involved in using azimuth monopulse to moreaccurately determine the cross-range location of a vessel. The measuredradial velocity field or color flow map has been utilized to rotate andorient the azimuth and elevation axes so that the center of the vesselis vertical, at approximately zero azimuth. The black circular cylinderrepresents the location of all points within the spatial resolution cellthat have a particular velocity.

FIG. 24 is a conceptual diagram illustrating the antifocusing techniqueaccording to an embodiment of the invention.

FIGS. 25 (A-F) illustrate a transmitter pattern wherein pulses of energyare transmitted according to an embodiment of the invention at sixdifferent ranges and center focused.

FIGS. 26 (A-F) show six different views of a combined transmit/receiveaccording to an embodiment of the invention.

FIGS. 27 (A-F) show Matlab simulations of transmitter patterns accordingto the invention, when continuous waves are utilized for purposes ofsimplification of illustration.

FIG. 28 is a 3-D illustration of the transmitter pattern illustrated inFIG. 27 (E).

FIG. 29 is a timing diagram illustrating the sequencing of transmitterbeam transmission according to an embodiment of the invention.

DETAILED DESCRIPTION OF THE INVENTION

The invention involves (1) a family of ultrasound sensors, (2) theinterplay of a set of core technologies that are unique by themselves,and (3) a number of design options which represent different ways toimplement the invention. To facilitate an organizational understandingof this many-faceted invention, a discussion of each of the three topicsabove follows.

The sensors addressed are all two-dimensional (i.e., planar or on thesurface of a convex shape such as a section of a cylinder) arrays ofpiezoelectric crystals for use in active, non-invasive, instantaneous(or real-time), three-dimensional imaging and monitoring of blood flow.The sensors use a unique approach to 3-D imaging of blood velocity andblood flow that (1) allows for finer image resolution than wouldotherwise be possible with the same hardware complexity (number of inputcables and associated electronics) and (2) allows for finer accuracythan would ordinarily be possible based on the resolution. The inventionmeasures and monitors 3-0 vector velocity rather than merely the radialcomponent of velocity.

Moreover, the present invention also utilizes (1) array thinning withlarge elements and limited scanning, (2) array shapes to reduce peaksidelobes and extend the field of coverage, (3) post-Dopplersub-resolution tracking, (4) post-Doppler sub-resolution mapping, (5)additional methods for maximizing the angular field of view, and (6)various digital beamforming procedures for implementing the mapping,tracking, and measurement processes. The present invention also extendsto array thinning, where the separation between array elements issignificantly larger than half the wavelength. This reduces the numberof input cables and input signals to be processed while maintaining highresolution and sensitivity and avoiding ambiguities. In a transcranialDoppler application, for example, where signal to noise and hencereceiver array area is of paramount importance, array thinning ispossible without reducing the receiver array area because a relativelysmall (compared to other applications) angular field of view is needed

Thinning with full aperture area imposes limitations on the angularfield of view Methods for expanding the field of view include using moreelements than are active at any one time—For example, if the electronicsare switched between two identical panels, the cross-range field of viewat any depth is increased by the size of the panel. If the panels arepointed in slightly different directions so that overlapping orredundant beams are avoided, the field of view is doubled. Ageneralization of this approach involves the use of an array on acylindrical or spherical surface.

Once a section of a blood vessel is resolved from other vessels inDoppler, depth, and two angles (az and el), Post-Doppler sub-resolutionprocessing locates that section to an accuracy that is one-tenth toone-twentieth of the resolution. This allows for precise tracking andaccurate mapping. Tracking provides for the possibility of unattendedlong term monitoring and mapping aids the operator in selecting thepoint or points to be monitored.

Furthermore, methods of the present invention permit non-invasive,continuous! 25 unattended, volumetric, blood vessel tracking, ultrasoundmonitoring and diagnostic device for blood flow. It will enableunattended and continuous blood velocity measurement and monitoring aswell as 3-dimensional vascular tracking and mapping using an easilyattached, electronically steered, transducer probe that can be in theform of a small pad for monitoring application, when desired. Moreover,a device and method of the present invention have applications inmeasuring the parameters described above in any part of the body. Anonlimiting example described below involves a cranial application.However as set forth, a device and method of the present haveapplications in any pad of the body, and can be used to track and mapany blood vessel in the body. A device of the present invention can, forexample:

-   -   1. Measure and continuously monitor blood velocity with a small        low-profile probe that can be adhered, lightly taped, strapped,        banded, or otherwise easily attached to the portion of the body        where the vascular diagnosis or monitoring required.    -   2. Track and maintain focus on multiple desired blood vessels in        spite of movement.    -   3. Map 3-D blood flow; e.g., in the Circle of Willis (the        central network of arteries that feeds the brain) or other        critical vessels in the cranial volume    -   4. Perform color velocity imaging and display a 3-D image of        blood flow that is rotated via track ball or joystick until a        desired view is selected.    -   5. Form and display a choice of projection, slice, or        perspective views, including (1) a projection on a depth-azimuth        plane, a B-scan, or a downward-looking perspective, (2) a        projection on an azimuth-elevation plane, a C-scan, or a        forward-looking perspective, or (3) a projection on an arbitrary        plane, an arbitrary slice, or an arbitrary perspective.    -   6. Use a track ball and buttons to position circle markers on        the points were measurement or monitoring of vector velocity is        desired.    -   7. Move the track location along the blood vessel by using the        track ball to slide the circle marker along the image of the        vessel.    -   8. Display actual instantaneous and/or average vector velocity        and/or estimated average volume flow.    -   9. Maintain a multi-day history and display average blood        velocity versus time for each monitored vessel over many hours.    -   10. Sound an alarm when maximum or minimum velocity is exceeded        or when emboli count is high; and maintain a log of emboli        detected.    -   11. Track, map, and monitor small vessels (e.g., 1 mm in        diameter), resolve vessels as close as 4 mm apart (for example),        and locate them with an accuracy of ±0.1 mm, for example.

Moreover, as explained herein, numerous methods have applications inobtaining the three dimensional coordinates of points along a bloodvessel from echoes returned from the body, and are encompassed by thepresent invention. A particular nonlimiting example of such a methodhaving applications herein is a novel and unobvious variation ofmonopulse tracking. For tracking purposes utilizing monopulse up to ninebeams are simultaneously formed for each transmit beam position. Inaddition to the “sum” beam that corresponds to the transmitted beam,there will either be 4 monopulse difference beams or there will be 8overlapping focused beams If a cluster of eight focused beams is used,these will be highly overlapped with the sum beam, and displaced veryslightly from the sum beam, with their centers equally spaced on a smallcircle around the center of the sum beam. These satellite beams wouldthen operate in pairs to form four difference beams. For example, theazimuth Monopulse ratio can be produced in two different ways, whichwill call “Liner” and “non-linear”. The non-linear method will determinethe magnitudes or the powers of three received signals, left, right, andsum (L. R, and S), and compute Ma=(|L|−|R|)/|S|. The linear method usescomplex signals and computes the azimuth monopulse ratio as the realpart of the ratio D_(a)/S₁ where D_(a)=L−R. D_(a) is the azimuthdifference.

For an ideal point target, the linear method for computing M_(a) resultsin an excellent estimate of the azimuth angle error. It also has theadvantage of only requiring 4, instead of 8 auxiliary beams. These 4beams would be an azimuth difference beam, D_(a), an elevationdifference beam, and two diagonal difference beams. The individualbeams, such as Land R, are not needed. However, beam shapes will behighly distorted by refraction through bone and tissue, and a“sub-optimum” non-linear approach might be more robust.

Regardless of which monopulse method is used, the conventional twodifference beams used in radar (azimuth difference and elevationdifference) may not be enough. The projection of the high-velocity dataon a plane perpendicular to the transducer line of sight (the C-scan)will usually be a line, not a point. With multiple difference beams,equally spaced in angle, one will be approximately perpendicular to theC-scan projection of the vessel. The system will select the monopulsedifference output with the largest magnitude. This provides anapproximate orientation of the C-scan projection of the vessel. Thecorresponding monopulse ratio (provided the sum beam power exceeds athreshold) is used to correctly re-steer and maintain a beam preciselycentered on that vessel.

If the power map output of a Wall filter is used for the monopulsebeams, the beam outputs are power and hence a complex ratio is notavailable. In that case the nonlinear method would be used. Analternative is to use the complex wall filter output, before computingthe power, with the linear method. During measurement, however, theoutput of a particular (high velocity) FFT Doppler bin may be used formonopulse (provided that the magnitude or power of the sum beam at thatDoppler exceeds a threshold). In that case either the linear or thenonlinear monopulse ratio may be used.

Another alternative is to use FFT processing and form the monopulseratio (linearly or non-linearly) at the output of a high-velocityDoppler-frequency cell with high sum-beam power. For example, set apower threshold and select the highest (positive or negative) velocitycell with power that exceeds the threshold. Since the data in a singleEFT cell is expected to be noisy, this procedure is recommended for ameasurement dwell, where enough time is spent in a single beam positionto have both useable velocity resolution and the ability to make severalmeasurements (multiple FFT's per frame).

FFT-Based Monopulse and Monopulse Averaging During Measurement

In a K pulse dwell, let K=K₁×K₂ where K₁ is the number of input pulsesused in the FFT and K₂ is the number of FFTs. Instead of performingmonopulse to re-steer the 20 beam every K₁ pulses, we compute themonopulse ratio at the output of a desired high velocity Doppler bin,and average its value over 1<2 FFTs. This reduces the steering noisewhile assuring that we are locating the center of the vessel (thehighest Doppler Energy). We chose the highest Doppler frequency forwhich the minimum sum beam power exceeds a threshold, and utilize onlythat Doppler cell for monopulse. The average is best performed as aweighted average. For example, if D_(n) and S_(n) are (say, elevation)difference-beam and sum beam outputs in the ntb FFT for the selectedDoppler bin, we chose:

${M = \frac{\sum\limits_{n = 1}^{K_{2}}{{{- S_{n}}}^{2}M_{n}}}{\sum\limits_{n = 1}^{K_{2}}{S_{n}}^{2}}},{{{where}\mspace{14mu} M_{n}} = {{Re}\left\{ {D_{n}/S_{n}} \right\}}}$${{or}\mspace{14mu} M_{n}} = \frac{{D_{n}}^{2}}{{S_{n}}^{2}}$

depending on whether linear or non-linear monopulse is used. For linearmonopulse it might be best to use only one large FFT (K₂=1). Fornon-linear monopulse, the expression simplifies to:

$M = \frac{\sum\limits_{n = 1}^{K_{2}}{D_{n}}^{2}}{\sum\limits_{n = 1}^{K_{2}}{S_{n}}^{2}}$

[Note that because a ratio is involved (so that beam pointing error isnot confused with signal strength) even the “linear method isnon-linear.]

A device of the present invention will allow a person with littletraining to apply the sensor and position it based on an easilyunderstood ultrasound image display. The unique sensor can continuouslymonitor artery blood velocity and volume flow for early detection ofcritical events. It will have an extremely low profile for easyattachment, and can track selected vessels; e.g., the middle cerebralartery (MCA), with no moving parts. If the sensor is pointed to thegeneral volume location of the desired blood vessel (e.g., within ±1cm.), it will lock to within ±0.1 mm of the point of maximum radialcomponent of blood flow and remain locked in spite of patient movement.

A device of the present invention can remain focused on the selectedblood vessels regardless of patient movement because it produces anddigitally analyzes, in real time, a 5-dimensional data base composed ofsignal-return amplitude as a function of:

1. Depth,

2. Azimuth,

3. Elevation,

4. Radial component of blood velocity,

5. Time.

Since a device of the present invention can automatically locate andlock onto the point with the maximum volume of blood having asignificant radial velocity, unattended continuous blood velocitymonitoring is one of its uses. By using the precise relative location ofthe point at which lock occurs as a function of depth, a device of thepresent invention can map the network of blood vessels as a3-dimensional track without the hardware and computational complexityrequired to form a conventional ultrasound image. Using the radialcomponent of velocity along with the three-dimensional blood path, adevice of the present invention can directly compute vector velocity.

A device used in a method of the present invention is a non-mechanicalDoppler ultrasound-imaging sensor comprising probes, processingelectronics, and display Specific choices of probes allow the system tobe used for transcranial Doppler (TCD), cardiac, dialysis, and otherapplications.

The present invention may be better understood by reference to thefollowing non-limiting Examples, which are provided as exemplary of theinvention. The following Examples are presented in order to more fullyillustrate particular embodiments of the invention. They should in noway be construed, however, as limiting the broad scope of the invention.

EXAMPLE 1 An Ultrasound Diagnostic and Monitoring Sensor with Real-Time3-D Mapping and Tracking of Blood Flow

This embodiment of the present invention has application for medicalevaluation and monitoring multiple locations in the body, however, thetranscranial Doppler application will be used as an example to describethe invention.

This invention provides: (1) affordable three-dimensional imaging ofblood flow using a low-profile easily-attached transducer pad, (2)real-time vector velocity, and (3) long-term unattendedDoppler-ultrasound monitoring in spite of motion of the patient or pad.None of these three features are possible with current ultrasoundequipment or technology.

The pad and associated processor collects and Doppler processesultrasound blood velocity data in a three dimensional region through theuse of a planar phased array of piezoelectric elements. Through use ofunique beamforming and tracking techniques, the invention locks onto andtracks the points in three-dimensional space that produce the locallymaximum blood velocity signals. The integrated coordinates of pointsacquired by the accurate tracking process is used to form athree-dimensional map of blood vessels and provide a display that can beused to select multiple points of interest for expanded data collectionand for long term continuous and unattended blood flow monitoring. Thethree dimensional map allows for the calculation of vector velocity frommeasured radial Doppler.

A thinned array (greater than half-wavelength element spacing of thetransducer array) is used to make the device inexpensive and allow thepad to have a low profile (fewer connecting cables for a given spatialresolution). The same physical array can also be used to form a broadtransmit beam encompassing a plurality of narrow receive beams. Initialacquisition of the blood velocity signal is attained by insonating alarge region by defocusing the transmit array or by using a smalltransmitting sub-aperture, for example. The computer simultaneouslyapplies numerous sets of delays and/or complex weights to the receiverelements in order to form M simultaneous beams. With M beams beingformed simultaneously, the receiver can dwell M times as long, so as toobtain high S/N and fine Doppler resolution. For an embodiment thatutilizes a small transmitting sub-aperture, the source of thetransmitted energy within the array (i.e., the location of thetransmitter sub-aperture) varies with time in order to lower thetemporal average spatial peak intensity to prevent skin heating.

The array is thinned without reducing the receiver area by limiting theangular field of view. When needed, a map of a larger region is createdby moving the probe and correlating the sub-images. Once the region ofinterest has been imaged and coarsely mapped, the full transmitter arrayis focused at a particular location on a particular blood vessel fortracking. In the tracking mode: (1) grating lobes due to array thinningare reduced by using wide bandwidth and time delay steering and (2) onlythree beams are formed at a time: sum, azimuth difference, and elevationdifference. When the desired (i.e. peak) blood velocity signal is not inthe output, this is instantly recognized (e.g., a monopulse ratio,formed after Doppler filtering, becomes non-zero) and the array is usedto track (slow movement) or re-acquire (fast movement) the desiredsignal. Re-acquisition is achieved by returning to step one to form andDoppler-process a plurality of beams in order to select the beam (andthe time delay or “range gate”) with the most high-Doppler (high bloodvelocity) energy. This is followed by post-Doppler monopulse tracking inazimuth, elevation, and range to lock a beam and range gate on to theexact location of the peak velocity signal.

In applications such as transcranial Doppler, where angular resolutionbased on wavelength and aperture size is inadequate, fine mapping isachieved, for example, by post-Doppler monopulse tracking each rangecell of each vessel, and recording the coordinates describing thelocation of the monopulse null. With a three-dimensional map available,true vector velocity can be computed. For accurate vector flowmeasurement, the monopulse difference is computed in a directionorthogonal to the vessel by digitally rotating until a line in theazimuth-elevation or C-scan display is parallel to the vessel beingmonitored.

All current ultrasound devices (including “Doppler color flow mapping”systems) form images that are limited by their resolution. In someapplications, such as TCD, the low frequency required for penetrationmakes the azimuth and elevation resolution at the depths of interestlarger than the vessel diameter. In this invention, as long as (1) ablood vessel or (2) a flow region of a given velocity can be resolved byfinding a 3-D resolution cell through which only a single vessel passes,that vessel or flow component can then be very accurately located withinthe cell. Monopulse is merely an example of one way to attain suchsub-resolution accuracy (SRA). SRA allows an extremely accurate map of3-D flow.

This invention utilizes post-Doppler, sub-resolution tracking andmapping; it does Doppler processing first and uses only highDoppler-frequency data. This results in extended targets since theactive vessels approximate “lines” as opposed to “points”. Inthree-dimensional space, these vessels are resolved, one from another.At a particular range, the azimuth-elevation axis can be rotated so thatthe “line” becomes a “point” in the azimuth dimension. That point canthen be located by using super-resolution techniques or by using asimple technique such as monopulse.

Overview of the Embodiment

The invention is complex because it involves (1) a family of ultrasoundsensors (for different parts of the body), (2) the interplay of a set ofcore technologies that are unique by themselves, and (3) a number ofdesign options which represent different ways to implement theinvention. To facilitate an organizational understanding of thismany-faceted invention, we precede a description of an overall preferredembodiment with a discussion of each of the three topics above.

The sensors addressed are all two-dimensional (i.e., planar) arrays ofin piezoelectric crystals for use in active, non-invasive, instantaneous(or real-time), three-dimensional imaging and monitoring of blood flow.While the sensors and the techniques for their use apply to all bloodvessels in the body, the figures and detailed description emphasizes thetranscranial Doppler (TCD) monitor because that application is mostdifficult to implement without all of the components of this invention.The sensors use a unique approach to 3-D imaging of blood velocity andblood flow that (1) allows for finer image resolution than wouldotherwise be possible with the same hardware complexity (number of inputcables and associated electronics) and (2) allows for finer accuracythan would ordinarily be possible based on the resolution. The inventionmeasures and monitors 3-D vector velocity rather than merely the radialcomponent of velocity.

The core technologies that constitute the invention are (1) arraythinning with suppression of ambiguities or grating lobes, (2)post-Doppler sub-resolution tracking, and (3) post-Dopplersub-resolution mapping. The invention encompasses two ways to thin thearray (reducing the number of input cables and input signals to beprocessed while maintaining high resolution and avoiding ambiguities).The first is bistatic operation; the second is broadband operation. Inthe TCD application, where signal to noise and hence receiver array areais of paramount importance, array thinning is possible without reducingthe receiver array area because a relatively small (compared to otherapplications) angular field of view is needed. One particular bistaticapproach to thinning reduces transmitter area and consequently poses aproblem of excessive spatial peak intensity (skin heating) in the TCDapplication. This is solved by a component invention called transmitterdiversity (which lowers the temporal average of the spatial peakintensity). The phase-defocusing bistatic approach and the monostatic orbistatic broadband approach to thinning all use the entire aperture andhence do not require transmitter diversity.

In the TCD application, the achievable angular resolution is poor,regardless of the method of thinning, or whether or not thinning isused. Once a section of a blood vessel is resolved from other vessels inDoppler, depth, and two angles (az and el), Post-Doppler sub-resolutionprocessing locates that section to an accuracy that is 10 to 20 times asfine as the resolution. This allows for precise tracking and accuratemapping. Tracking provides for the possibility of unattended long termmonitoring and mapping aids the operator in selecting the point orpoints to be monitored.

There are many options available in the design of any member of thefamily of sensors that utilizes any or all of the core technologies thatcomprise this invention. A two-dimensional array is established art thatcan be designed in many ways and can have many sizes and shapes(rectangular, round, etc.). Digital beamforming (DBF) is a techniquethat has been in the engineering literature (especially radar and sonar)for many years. One medical ultrasound DBF patent cites many references,while another describes a particular instance of DBF without citing theother patent or any other prior art. While planar arrays, DBF, Dopplerultrasound, and color flow imaging are prior art, the manner in thisspecification of using such established technologies to map, track,measure, and monitor blood flow is unique. The embodiment is anon-invasive, continuous, unattended, volumetric, blood vessel tracking,ultrasound monitoring and diagnostic device. It will enable unattendedand continuous blood velocity measurement and monitoring as well as3-dimensional vascular tracking and mapping using an easily attached,electronically steered, transducer probe that can be in the form of asmall pad for monitoring application, when desired. Although the devicehas application to multiple body parts, the cranial application will beused as a specific example. The device can, for example:

-   -   1. Measure and continuously monitor blood velocity with a small        low-profile probe that can be adhered, lightly taped, strapped,        banded, or otherwise easily attached to the portion of the body        where the vascular diagnosis or monitoring is required.    -   2. Track and maintain focus on up to four desired blood vessels        in spite of movement.    -   3. Map 3-D blood flow; e.g., in the Circle of Willis (the        central network of arteries that feeds the brain).    -   4. Perform color velocity imaging and display a 3-0 image of        blood flow that is rotated via track ball or joystick until a        desired view is selected.    -   5. Form and display a choice of projection, slice, or        perspective views, including (1) a projection on a depth-azimuth        plane, a B-scan, or a downward-looking perspective, (2) a        projection on an azimuth-elevation plane, a C-scan, or a        forward-looking perspective, or (3) a projection on an arbitrary        plane, an arbitrary slice, or an arbitrary perspective.    -   6. Use a track ball and buttons to position circle markers on        the points at which we wish to measure and monitor vector        velocity.    -   7. Move the spatial resolution cell being measured along the        blood vessel by using the track ball to slide the circle marker        along the image of the vessel.    -   8. Display actual instantaneous and/or average vector velocity        and/or estimated average volume flow.    -   9. Maintain a 3-day history and display average blood velocity        versus time for each monitored vessel over 14 hours.    -   10. Sound an alarm when maximum or minimum velocity is exceeded        or when emboli count is high.    -   11. Track, map, and monitor vessels as small as 1 mm in        diameter, resolve vessels as close as 4 mm apart (for example),        and locate them with an accuracy of +0.1 mm.

The Monitoring Device will allow a person with little training to applythe sensor and position it based on an easily understood ultrasoundimage display. The unique sensor can continuously monitor artery bloodvelocity and volume flow for early detection of critical events. ft willhave an extremely low profile for easy attachment, and can trackselected vessels; e.g., the middle cerebral artery (MCA), with no movingparts. If the sensor is painted to the general volume location of thedesired artery (e.g., within ±C.5 cm.), it will lock to within ±0.1 mmof the point of maximum radial blood flow and remain locked in spite ofpatient movement.

The device can remain focused on the selected blood vessels regardlessof patient movement because it produces and digitally analyzes, in realtime, a 5-dimensional data base composed of signal-return amplitude as afunction of:

6. Depth, 2. Azimuth, 3. Elevation, 4. Radial blood velocity, 5. Time.

Since the device can automatically locate and lock onto the point withthe maximum volume of blood having a significant radial velocity,unattended continuous blood velocity monitoring is one of its uses. Byusing the precise relative location of the point at which lock occurs asa function of depth, the device can map the network of blood vessels asa 3-dimensional track without the hardware and computational complexityrequired to form a conventional ultrasound image. Using radial velocityalong with the three-dimensional blood path, the device can directlycompute vector velocity.

The proposed device is a non-mechanical Doppler ultrasound-imagingsensor consisting of probes, processing electronics, and display.Specific choices of probes allow the system to be used for transcranialDoppler (TCD), cardiac, dialysis, and other applications.

FIG. 1 shows the TCD configuration and the initial definition of thedisplay screen. The TOO system is comprised of one or two probesattached to the head with a “telephone operator's band” or a Velcrostrap. The interface and processing electronics is contained within asmall sized computer. A thin cable containing 64 micro coax cablesattaches the probe to the electronics in the computer. When the operatorpositions the probe on the head the Anterior, Middle and PosteriorCerebral Arteries and the Circle of Willis are imaged on the screenalong with other blood vessels. The arteries or vessels of interest areselected by viewing the image. The system locks onto the blood vesselsand tracks their position electronically. A variety of selectedparameters is presented on the screen; e.g., the velocity, the pulserate, depth of region imaged, gain and power level. Using only one probethe TCD can monitor up to two arteries (vessels) at a time. Presented onthe screen are dual traces, one for each artery. The blood velocity canbe dynamically monitored. As shown in FIG. 1 both the current bloodvelocity (dark traces) and any historic trace (lighter color) can bedisplayed simultaneously. The average blood velocity or estimatedaverage flow for each artery is displayed below the respective velocitytrace. The image shows the arteries and the channel used for eachartery. When two probes are used, the display is split showing signalsfrom both of them. Using a different probe (i.e., different size) withthe same electronics and display, the unit can be used to measure andmonitor the blood flow in a carotid artery. Similarly, it can be used toperform this function for dialysis, anesthesia, and in other procedures.

The sensor is a two dimensional array of transducer elements(piezoelectric crystals) that are configured and utilized differentlyfor transmit and receive during acquisition. For example, if a square(N×N) array is used, all N² elements would receive at the same time, butonly a 2×2 sub-aperture would transmit at any one time. This isillustrated in FIG. 2 for the case of N=8. The array need not be square.Any M×N array may be utilized in this manner. All NM received signals(64 in our example) are sampled, digitized, and processed. This can bedone, for example, in a desk top or lap top personal computer withadditional cards for electronics and real-time signal processing asillustrated in FIG. 1 and FIG. 3. If the PCI bus in FIG. 2 becomes abottleneck for high speed processing, a pipelined or systolicarchitecture would be used. Alternatively, the processing can beperformed in an application specific integrated circuit (ASIC).

The small (4 element) transmit sub-aperture (FIG. 2 b) produces a broadtransmit beam that insonates a region containing many receive beams.This is schematically illustrated in FIG. 4 for the particular case of asquare array and square elements such as in FIG. 2. Since data isreceived from each element of the array, this data can be combined in aprocessor (FIG. 3, for example) in many different ways to form anynumber of beams. The transmitter is larger than a single array elementso that it can provide some selectivity and not insonate the gratinglobes caused by array thinning (spacing the array elements more than ½wavelength apart). The concept is illustrated below for a 1-dimensionalarray forming a beam that measures only one angle. For a two-dimensionalarray, this represents a horizontal or vertical cut through the clusterof beams shown in FIG. 4. FIG. 4 was an approximate and conceptualrepresentation of the two-angle (azimuth and elevation) extension of thesingle angle case detailed below.

“Grating lobes” are ambiguities or extra, unwanted, beans caused byusing a transducer array whose elements are too large and hence too farapart. The following analysis illustrates grating lobe suppression forthe worst case of narrowband signals and phase-shift beam processing.Time delay processing using wideband signals would be similar, but wouldfurther attenuate or eliminate grating lobes, resulting in even betterperformance.

The next four figures show beam pattern amplitudes plotted against

x=(d/λ)sin θ,  (1)

where x represents a normalization for the angle, θ, from whichreflected acoustic energy arrives. The azimuth (or elevation) angle, θ,is zero in the broadside direction, perpendicular to the transducerarray. The width (or length) of a transmitter is 2d, where d is thewidth (or length) of a single element of the receiver array. Thewavelength of the radiated acoustic wave is λ=c/f where c is theacoustic propagation velocity (1543 meters/second in soft tissue) and fis the acoustic frequency (usually between 1 and 10 megahertz). FIG. 5 ashows the transmitter pattern

a _(T)(x)=sin 2πx/2πx  (2)

for the special case of uniform insonation over the 2d-wide transmittersub-aperture being used.

The receiver pattern is the product of the receiver element pattern andthe receiver array pattern

a _(R)(x)=a _(RE)(x)a _(RA)(x)  (3)

Each of these two component patterns is plotted separately in FIG. 5 b.Again assuming the special case of a uniform receiver element (and asquare element in the case of a 2-D array), the receive element patternis

a _(RE)(x)=sin πx/πx  (4)

The receiver element pattern is twice as wide as the transmitter patternbecause the receiver element is half as wide as the transmitter. In thefar-field, i.e., for λr>>L², where r is the range or depth and L is thelength of the aperture, the receive array pattern steered to the angleθ=θ₀ is

$\begin{matrix}{{{a_{RA}(x)} = {\sum\limits_{n = 0}^{N - 1}{w_{n}^{{j2\pi}\; {n{({x - x_{0}})}}}}}},} & (5)\end{matrix}$

where w_(n) is a weighting to reduce sidelobes and N is the number ofelements in one dimension. As seen in FIG. 5 b, equation (5) is periodicin x. The peak at x=x₀(x₀=0 in FIG. 5) is the desired beam and theothers are grating lobes.

In the near field, when focused at (r₀, θ₀ equation (5) is replaced bythe slightly better general Fresnel approximation:

$\begin{matrix}{{a_{RA}\left( {x,z} \right)} = {\sum\limits_{n = 0}^{N - 1}{w_{n}^{{j2\pi}{\lbrack{{n{({x - x_{0}})}} + {{({n\; \frac{N - 1}{2}})}^{2}{({z - z_{0}})}}}\rbrack}}}}} & (6)\end{matrix}$

(provided that that the range significantly exceeds the array size,r>L). where x=d sin θ/λ, as before, and

z=d ² cos² θ/λr  (7)

Because the receiver aperture is sampled with a spatial period of d, thereceiver array pattern will be periodic in sin θ, with a period of λ/θ(equation 5) This periodicity means that the array pattern is ambiguous.When the array is pointed broadside (θ=0), it will also be pointed atthe angle θ=sin⁻¹ (2/d), for example. In terms of the normalizedvariable, x, the period is unity. Since I sin θ| cannot exceed 1, thevariable x is confined to the interval [−d/λ, d/λ]. The conventionalelement spacing is d=λ/2. Thus, in a conventional phased array, x isalways between −0.5 and +0.5, and hence ambiguities are not encountered.In a highly thinned array (d>2) there will normally be ambiguities orgrating lobes as illustrated in FIG. 5 b. The second grating lobe, atx=2 or 0=sin⁻¹ (2λ/d), is not real when d does not exceed 2λ.

FIG. 5 c shows the two-way pattern. The gating lobe suppression,resulting from the choice of a transmitter diameter of D=2d is valid foralt values of d. In a two dimensional array, the elements could berectangular instead of square (d_(x)×d_(y)), and the results would stillbe valid. Similar results could be obtained for an array in which theelements are staggered from row to row (and/or column to column). Forexample, if the receiver array is a “bathroom tile” of hexagonalelements, the transmitters could be chosen as sub-arrays consisting ofan element and its six surrounding neighbors.

In FIG. 6 the same array is used as in FIG. 5, but the receiver elementsignals are combined with a phase taper that steers the beam to x=0.2.This is approximately (a little less than) the half power point, wherea₁(x) a_(re)(x)=0.707. In FIG. 6 c, we see that the grating lobes arenot completely suppressed, with the largest one at x=−1+0.2=−0.8. FIG. 7shows this in decibels. The worst-case grating lobe is attenuated by atleast 25 dB, even in the stressing case of extremely narrow bandoperation. A Hanning window was applied to keep the sidelobes lower thanthe peak grating lobe. These Figures were produced in MATLAB, using thefollowing software (m-file):

-   -   x=−2:1/64:2−1/64;        p=pi*x+eps; R=sin(p)·/p;        p=2*p; T=sin(p)·/p;

N=8

n=0:N−1;% xo=0;xo=0.2; % is 2-way ½ powere=exp (j*n′*2*pi(x−xo));w=hanning(N);

% E=(1/N)*ones(1, N)*e; E=(2/N)*w′*e;

subplot(311); plot (x,abs (T));subplot(312); plot (x, [abs (R);abs (E)]); ITRE=abs(T).*abs(R).abs(E);subplot(313); plot(x,TRE);FIG. (2); plot(x,20*log 10(TRE));zoom on;

The dimensions in FIG. 4 are representative for a transcranial Dopplerapplication of the invention, to provide a specific example. If f=2 MHzis chosen for the center frequency, the wavelength is 0.77 mm. An 8×8array with a width and/or length of L=1 cm, provides a one dimensionalthinning ratio of 2 d/λ=3.247. For a square array, the total number ofelements is reduced by a factor of (2 d/λ)²≧10 from that of a filledarray. Even greater thinning ratios are possible. Even if d/λ is keptless than 2 to avoid a second grating lobe (at x=2), complexityreductions up to a factor of 16 are possible. For the 1 cm array at 2MHz, the hyperfocal distance (where the 3 dB focal region extends toinfinity) is L²/4λ=3.25 cm. Thus, a fixed focus probe suffices for thisapplication. However, since the simultaneous formation of multiplereceive beams is conveniently performed digitally, dynamic focus onreceive is easily accomplished. The quadratic phase distribution acrossthe elements required to focus in depth is simply added to the linearphase distributions required to steer the beams.

FIG. 8 a shows the product of the transmitter pattern (FIG. 5 a or 6 a)and the receiver element pattern. FIG. 8 b plots the element patternsfor a set of five beams steered to x=−0.2, −0.1, 0, 0.1, and 0.2. Thisset of five receive beams shows grating lobes of the thinned array. FIG.Sc shows the set of resulting 2-way patterns obtained by multiplying thepatterns in FIG. 8 b by the function plotted in FIG. 8 a. Here, thegrating lobes are suppressed. This represents a horizontal or verticalcut through the cluster of beams in FIG. 4.

Using the configuration described above, the cluster of beams in FIGS. 4and 8 c is used to approximately locate the desired point for collectingthe blood velocity signal. For example the output of each beam in thecluster would be Doppler processed by performing an FFT or equivalenttransformation on a sequence of pulse returns. The pulse repetitionfrequency (PRF) would typically be less than or equal to 9 kHz tounambiguously achieve a depth of 8.5 cm for the TCD application. Inorder to obtain a velocity resolution as fine as Δv=1 cm per second (todistinguish brain death), a dwell of duration T=λ/(2Δv)=38.5 ms, or 347pulses at 9 kHz, is desired. For efficient FFT processing, the number ofpulses used would be zero filled to a power of 2 such as 512.

The example shown in FIGS. 2 through 8 was an 8 by 8 receiver arrayforming a 5 by 5 cluster of beams. This is an example of an approximaterule of thumb for this invention, that an N element linear array isrecommended for use in producing N/2+1 beams for N even and [N+1]/2beams for N odd. Thus, a 16 by 10 element rectangular array wouldpreferably be used to form a 9 by 6 cluster of beams, though the actualnumber of beams formed is arbitrary. This recommended number of beams isderived below. If an N elements were used to form orthogonal beams,e.g., by an N-point FFT, then there would be N beams in a 180° angularregion, from −90° to +90°, corresponding to −1≦u≦1, where u=sin θ. Inconventional phased array ultrasound, a 128 (=N) element array is usedto produce 256 (=2N) lines (sequentially scanned beams) in a 90° angularregion from −45° to 45°, corresponding to −0.707≦u≦0.707. If the arrayis filled, then x=u/2 (Equation 1) and 2N beams are conventionallyformed in |x|<√2/4. When we thin the array, we prefer to have|x|<0.2=1/5 (the 3 dB point of the curve in FIG. 7 a). The number ofbeams in that region, for the same beam density as used in currentpractice, is given by:

Recommended No. of beams=(1/5)N÷(√2/4)=2√2N/5≈0.5657N.

The beams are formed digitally, using software on a personal computer orusing digital signal processing hardware to implement equations such asEquation 5 or 6. The electronic interface between the probe and theprocessor is diagrammed n FIG. 9. This figure illustrates the case ofsignals from 64 elements being connected to a single A/D converter, andpower being applied to sets of four elements. The use of a separate A/Dconverter for every received channel, for example, is another possibleimplementation of this invention.

A conventional, half-wavelength spaced, monostatic, phased array could25 sequentially search a region of interest, butt would require far moreelements and would thus be far more costly. Using the array differentlyin transmit and receive, not only allows for the formation of multiplebeams; it also enables the use of the angular pattern of the transmitterto suppress receiver grating lobes. This allows for a “thinned” array(elements spaced less than a half wavelength apart). Because receivebeams are formed only in a limited angular region, a wide-angle receiverelement pattern (which usually implies a small element) is not required.In fact, the size of the receiver element can be as large as the elementspacing. Thus the receiver array is “thinned” only in the sense that theelement spacing exceeds a half wavelength. Since the element size alsoexceeds a half wavelength, the array area is not reduced. It is thinnedonly in terms of number of elements, not in terms of receiver area.Consequently, there is no reduction in signal-to-noise ratio, for arequirement for increased transmitter power.

A monostatic array would transmit from the full aperture, scanning thetransmitted beam over the region being examined. The “bistatic” array ofthis invention transmits from a sub-aperture to insonate multiplereceive beam positions simultaneously. Since there is an FDA limit tospatial peak, temporal average, intensity (I_(spta)) there may be adanger of exceeding this limit at the transducer surface, creating adanger of burning the skin. This potential danger is eliminated by usinga different transmit sub-aperture for each coherent dwell or burst ofpulses. This transmitter diversity technique spreads the temporalaverage intensity over the face of the array, reducing I_(spta) to whatit would be if the entire array were used at once.

For the particular implementation pictured in FIG. 9, an A/D converteris multiplexed amongst the 64 elements. The signal spectrum at any ofthese elements is centered at f₀=2 MHz, as shown in FIG. 10 a. This is areal signal with a spectrum that is symmetric about f=0. This analogsignal is bandpass filtered (BPF) to insure that there is little poweroutside of a 444 kHz band centered at 2 MHz. If a 512/9 56.889 MHz A/Dconverter is used, each receive channel is sampled f_(s)=888.9 kHz,giving rise to a real sampled signal with a spectrum as shown in FIG. 10b. A processing element such as a field programmable gate array (FPGA)is used to shift the frequency by f_(s)/4 (FIG. 10 c) by “multiplying”by quarter cycle samples of sinusoids (which are zeros and ones). Thesame FPGA also digitally filters (or Hubert transforms) the complexsignal to decimate its sampling rate by a factor of two. The spectrum ofthe decimated digital low-pass signal is shown in FIG. 10 d.

The signal sent to the processor from each element has the spectrumshown in FIG. 9 d, and consists of r=f_(s)/2 complex samples per second.The total data rate into the processor is approximately 57 megabytes persecond. For non-real-time operation, tens of seconds of data at a timewill be collected in system memory and then transferred to hard disk.For real-time monopulse tracking, only three beams are formed, so thatthe data rate is reduced to 3×0.8889=2.67 Mbytes, or 5.33 Mbytesallowing for bit growth.

The transmitted pulses are sent to a group of four elements Theparticular embodiment shown in FIG. 9 uses diodes to block the receivedsignals and prevent mutual coupling between the four receive elements.After a coherent pulse train (or pulse burst used for Dopplerprocessing), the waveform is switched to another set of 4 elements forthe next burst. A separate power amplifier is associated with each ofthe 16 sets of elements so that the switching can be accomplished at lowpower.

One embodiment of sub-resolution tracking (i.e., tracking and locatingblood flow to a small fraction of a spatial resolution cell) is“Monopulse”. Monopulse tracking is performed as follows. A particularset of complex weights are applied to the set of received signals (64 inthe example of FIG. 2) to steer a beam at the middle cerebral artery,for example. The phase taper across the array defines the steeringdirection and the amplitude taper (called a window in radar and ashading in sonar) is used to provide low sidelobes for high dynamicrange. The beam output (a linear combination of the signals) is rangegated (time delay corresponding to the desired depth) and therange-gated/beam-formed output from a sequence of transmitted pulses isthen Fourier transformed to obtain a plot of amplitude versus Dopplerfrequency. The receive beam is steered digitally to the point thatproduces the maximum amplitude at high Doppler frequencies.

Since the measured data at each element is stored, the digital processorcan apply more than one set of weights at a time, forming more than onebeam. For software monopulse the processor will form three beams, all inthe same direction. All three beams may have the same phases applied tothe element signals; but the amplitudes will differ. The beam called Sumhas all positive amplitudes, with the larger weights applied to thecentral elements. This forms a fairly broad beam. The beam called A_(z)for “azimuth difference beam” has large positive weights on therightmost elements and large negative weights on the leftmost elements(or vice versa). The beam called El for elevation difference has largepositive weights on the top-most elements and large negative weights onthe bottom-most elements. A correctly pointed beam would haveA_(z)=El=0, and Sum would be maximized.

The ratio of the peak Doppler amplitude outputs: Az/Sum, is a precisemeasure of the azimuth pointing error and the corresponding ratio El/Summeasures the elevation pointing error. The digital steering phase taperis thus corrected with data from a single burst of pulses. The durationof the pulse burst is the reciprocal of the medically required Dopplerresolution (usually corresponding to the minimum blood velocity that cansupport life). Without techniques such as those described in thisspecification, a sequence of at least four additional Doppler dwells orpulse bursts would be required (above, below, to the right, and to theleft) in a hunt and seek method to find the correct (maximum peakDoppler Amplitude) beam. With monopulse, the correction is very precise(to within ±0.1 mm of the point of maximum peak Doppler amplitude) andvirtually instantaneous. For the bistatic digitally beamformed sensor,the original data exists in computer memory. Hence, whenever the Dopplerprocessed monopulse differences are non zero, the same data set couldeven be re-processed to form a correctly pointed beam. A slowerprocessor would merely process the next burst correctly.

A “front view” perspective display or a C scan display (azimuthhorizontal and elevation vertical) of the blood flow map at the desiredrange will allow someone to aim the transducer probe or pad at thedesired point (highest amplitude for high Doppler), so that the desiredpoint is initially within the center beam. The receiver array is thensteered electronically so that the monopulse differences are zero andhence the central beam is precisely aimed at the desired point. Slightmotions are corrected using monopulse and large motions are corrected byagain forming alt beams to re-acquire the peak signal. All correctionsare made entirely electronically, in the data processing or digitalbeamforming. A narrow receiver beam will always be precisely pointed (towithin a tenth or 20^(th) of the receiver beamwidth) as long as thedesired point remains within the much larger region covered by thetransmitter (FIG. 4).

True vector velocity is computed from the blood vessel map and theradial velocity measured from the pulse Doppler dwell. A map, far moreaccurate than that attainable with the available angular resolution isattained as follows. The low-resolution map is used to locate a vesselof interest and a beam is locked on it at a fixed range, using azimuthand elevation monopulse. The coordinates of the point at which lockoccurs is recorded. The range is then changed slightly, another lock (onthe same vessel) is obtained, and the coordinates are recorded. In thismanner, the vessel is mapped far more accurately than would be predictedfrom the available image resolution. All vessels within the field ofview of the probe are similarly mapped. By moving the probe angleslightly, another region can be mapped in the same manner. Several suchmaps can be correlated over the region of pair-wise overlap andconverted to a common coordinate system. In this manner a larger regionis mapped and displayed than that of the current field of view. Thecurrent field of view would be highlighted, outlined, or presented as acolor flow map. Points to be monitored in the current field are thenselected by moving a cursor along the display (point and click). Theselected points are Doppler processed and tracked usingthree-dimensional monopulse. While Doppler measurements provide only theradial component of velocity, the accurate blood vessel map provides theexact three-dimensional orientation of the vessel at the point beingmonitored. The measured radial velocity is divided by the projection ofa unit vector representing the vessel at the monitored point onto thetransducer line of sight. This gives the magnitude of the true vectorblood velocity.

Sub-resolution mapping accuracy is attainable if (1) therange-azimuth-elevation-Doppler resolution cell being examinedencompasses only a single blood vessel, and (2) “azimuth” monopulse isperformed with the usually vertical e-axis tilted so that theorientation of the vessel in the spatial resolution cell being processedis parallel to the e-r plane (“azimuth” is constant).

The user will ascertain from the display, that the resolution cell beingmonitored contains only a single vessel, and would rotate the 3-Dblood-vessel map to a C-scan aspect (elevation up and azimuth to theright). A vertical mark will appear in the display, within theresolution circle, to signify the orientation of the monopulse axis.This axis (parallel to the line separating the positively and negativelyweighted array elements) can then be oriented so that the mark isaligned with the blood vessel in either of two ways. The probe can bephysically twisted (rotated about the line of sight), or it can beelectronically rotated via digital processing because the weights areapplied digitally.

FIG. 11 illustrates the segment of a vessel in a single resolution cell,after rotation. The resolution cell shown is not a cube because therange resolution will typically be finer than the cross-rangeresolution. The illustrated circular cylinder represents blood cells ina vessel reflecting energy at a fixed Doppler frequency. These representa cylindrical annulus of blood cells, at a constant distance from thevessel wall, moving with approximately the same velocity. In the singleresolution cell of FIG. 11, the return at the highest Doppler wouldrepresent a line in three-dimensional space (the axis of the vessel) andhence a point on the azimuth axis after rotation. When applied to thehighest Doppler output, the Sum beam would have broad peak at zeroazimuth (a=0) and the monopulse ratio, r=Az/Sum, will be a linearfunction of the azimuth angle to which the array is phase steered:

r(a)=ka.

This result can be attained by applying the same phase across theaperture for the Az and Sum beams, but using the derivative of the Sumbeam amplitude weights with respect to x and y respectively for the Azand El aperture weights.

Other Embodiments

If the wide transmit beam (for search and acquisition) is created byusing a quadratic phase curvature instead the scheme of FIG. 2 b,transmitter diversity may not be needed. Furthermore the manner ofcontrolling grating lobes in FIG. 1 and FIGS. 5-8 is only one of many.Using a wider bandwidth and time-delay steering can also reduce gratinglobes.

EXAMPLE II Ultrasound Measurement of Blood Volume Flow

As explained above, current ultrasound Doppler devices measure radialvelocity. Several methods now exist for 3-D ultrasound imaging, such asthose involving transducer motion. A three-dimensional image withDoppler allows for the measurement of vector velocity. Example I aboveprovides measurement and long term monitoring of three-dimensionalvector velocity. If the resolution of a color flow Doppler image issufficient to provide an estimate of the inside diameter of the bloodvessel, then measurement of volume blood flow becomes practical.Presently available ultrasound imaging devices have either lowresolution or they only produce a two-dimensional image. The presentinvention combines vector velocity information (such as attained asexplained in Example I above) with additional information to obtainvolume flow. The additional information is the inside diameter of thevessel under examination, the blood velocity profile across the vessel,or the vector velocity as a function of time and position (i.e., thevelocity field). This additional information can be obtained from ahigh-resolution radial-Doppler or color flow image or from external datasuch as a high-resolution MRI image.

A two-dimensional array of piezoelectric elements, or some other means,is used to image blood flow in a three dimensional region. A particularpoint on a particular vessel is selected and the vector representing theorientation of the vessel is noted. The radial velocity divided by thecosine of the angle made by the vessel with the line of sight at themeasurement point is the magnitude of the vector velocity. That numberintegrated over the vessel cross section would give the volume flow involume per unit time or milliliters per minute, for example.

FIG. 12 shows a circular cylinder representing blood cells in a vesselmoving at a particular velocity and thus reflecting energy at a specificDoppler frequency. The figure assumes that methods such as those in thereferenced invention, for example, have been used to measure the 3-Dorientation of the vessel so that the vector velocity can be calculatedand the azimuth axis can be defined to be perpendicular to the vessel.

The simplest way to estimate volume flow is to measure the vesseldiameter, d, (or radius d/2), calculate the cross-sectional area,A=π(d/2)², and multiply by the average velocity. A more accurate way isto integrate the velocity as a function of position, over thecross-section. The velocity is a function of the radius, a. of thecylinder depicted in FIG. 12. If a is the distance from the cylinder toits axis, and v (a) is the velocity function, then the volume flow is

$\begin{matrix}{2\pi {\int_{0}^{d/2}{{{av}(a)}\ {a}}}} & (7)\end{matrix}$

Equation (1) assumes a circular cross-section of constant radius r=d/2.It is a special case of the more general polar coordinate integration:

$\begin{matrix}{\int_{0}^{2\pi}{\left( {\int_{0}^{r{(\theta)}}{{{av}\left( {a,\theta} \right)}\ {a}}} \right)\ {\theta}}} & (8)\end{matrix}$

The velocity function is determined by determining the diameter (andhence the radius) of the cylinder corresponding to each velocity.

For example a 1.5-cm diameter Doppler ultrasound transducer arrayoperating at 10 MHz will be oriented with the length or azimuthdirection perpendicular to the vessel to produce a B-scan(depth-azimuth) image. At a depth of approximately 10 mm, the crossrange resolution is 0.1 mm. If the vessel diameter is 1 mm, the diametercan be measured with an accuracy of +5%. The area of the vessel is thusknown to an accuracy of 10%. Since the average vector velocity can bemeasured extremely accurately, the volume flow is also accurate to +10%.The best accuracy is attained by measuring the azimuth extentcorresponding to various velocities and then numerically evaluatingequation (7) or (8). Naturally, a skilled artisan can readily program aprocessor to solve these equations, and calculate blood flow volumeusing routine programming techniques.

Since the autocorrelation function (pulse-to-pulse, at a fixed range)and the Doppler Power Spectrum form a Fourier pair, the total power canbe obtained either as the autocorrelation function at zero Lag or theintegral of the Doppler Power Spectrum (Spectral Density) over allDoppler frequencies. Since radial velocity is proportional to Dopplerfrequency, the mean velocity can be obtained from the autocorrelationfunction as shown below:

${R_{xx}(\tau)} = {{\frac{1}{2\pi}{\int_{- \infty}^{\infty}{{S(\omega)}^{j\omega\tau}\ {\omega}}}} = {\int_{- \infty}^{\infty}{{S_{d}(f)}^{{j2\pi}\; f\; \tau}\ {f}}}}$henceR_(xx)(0) = ∫_(−∞)^(∞)S_(d)(f) f = P_(d) = total  Doppler  powerand${{R_{xx}(\tau)} = {{R_{xx}(\tau)} = {\int_{- \infty}^{\infty}{\left\lbrack {j\; 2\pi \; {{fS}_{d}(f)}} \right\rbrack ^{{j2\pi}\; f\; \tau}\ {f}}}}},{{{leading}\mspace{14mu} {to}{R_{xx}(0)}} = {{{j{\int_{- \infty}^{\infty}{\left\lbrack {2\pi \; {{fS}_{d}(f)}} \right\rbrack ^{{j2\pi}\; f\; \tau}\ {{f}.{Thus}}}}} - {j\; \frac{R_{xx}(0)}{R_{xx}(0)}}} = {{2\pi {\int_{- \infty}^{\infty}{f\; \frac{S_{d}(f)}{\int_{- \infty}^{\infty}{{S_{d}\ (f)}{f}}}\ {f}}}} = {2\pi \; E\left\{ f_{d} \right\}}}}},$

where the Doppler frequency and its mean (expected value) arc related tothe radial blood velocity and its mean by

$f_{d} = {\frac{2f_{0}}{c}{v.{Hence}}}$${E\left\{ v \right\}} = {{\int_{- \infty}^{\infty}{v\; \frac{P(v)}{\int_{- \infty}^{\infty}{{P(v)}\ {v}}}\ {v}}} = {{- j}\; \frac{c}{4\pi \; f_{0}}\frac{R_{xx}\left( 0_{5} \right)}{R_{xx}(0)}}}$

which is used in the autocorrelation method of color-flow blood-velocityimaging. We note that if we do not normalize by dividing by the totalDoppler power, we obtain a power-velocity product that indicates thevolume flow rate. This is due to the fact that power is directlyproportional to area [see Reference 1].

Since all velocity vectors are parallel at the narrowest point (the venacontracta), flow at that particular point can be considered asnon-turbulent, even though severe turbulence exists before and after.Reference I shows that regurgitant blood flow through the mitral heartvalve can be quantitatively measured by observing the Doppler spectrumat that point and using the power-velocity-integral relation below

In terms of the velocity power spectrum, P(v)=(2f₀/c) S_(d)(f) Reference1 shows that the blood vessel area in a “slice” perpendicular to theline of sight is directly proportional to the total Doppler power (thetotal power at the output of the high-pass 5 wall filter).

$A = {{\frac{A_{0}}{P_{0}}P_{d}} = {\frac{A_{0}}{P_{0}}{\int{{P(v)}{v}}}}}$

where A₀ and P₀ are the known area and measured power in a narrow beam,smaller than the vessel. If the blood flow makes an angle θ with theline of sight, the area, and hence power, is increased by the factor1/cos θ. This offsets the fact that only the radial component ofvelocity is measured, so that the power velocity integral provides truevolume flow:

$Q = {{{Q}/{t}} = {\frac{A_{0}}{P_{0}}{\int{{{vP}(v)}{v}}}}}$

In Reference 1, P was measured with the same probe as P₀ by masking theoutside of the aperture in order to create a wider beam. With our 2-Dphased array, we would merely turn off or ignore some of the outerelements. More importantly, we can use the 3-D image to precisely locatethe vena contracta, and lock onto it using monopulse. We can evenmonitor the valve during a stress test, while the patient is on atreadmill.

We note here, that there are several ways to measure the volume flowrate. Reference 1 uses the fact that it is proportional to the integralof the product of the velocity and the power per unit velocity, as inthe last equation. Another way is to recognize that it is equal to theproduct of the average radial velocity and the total projected area thatis, in turn, proportional to total Doppler power. Since the totalDoppler power is used in the denominator of the autocorrelation-basedcolor-flow velocity map, volume flow rate can be obtained by merely notdividing by the total power. If the i^(th) pulse return (after MTI orDoppler high-pass or wall filtering) is

z _(i) =x _(i) +jy _(i) i=1,2, . . . N,

the volume flow rate is proportional to

${\sum\limits_{i = 2}^{N}{x_{i}y_{i - 1}}} - {y_{i}x_{i - 1}}$

The normalization (denominator) that is used to convert this lastquantity to mean velocity can be

${\sum\limits_{i = 2}^{N}{x_{i}x_{i - 1}}} + {y_{i}y_{i - 1}}$

that is based on a derivation in Reference 2, or a simple powerestimate, such as

${\sum\limits_{i}x_{i}^{2}} + y_{i}^{2}$

The point we wish to make here is that by not dividing by a powerestimate to obtain radial velocity, we obtain volume flow. Currentultrasound Doppler imaging systems compute the mean velocity as a ratio,

E(v)=F/P _(d,)

and display it as a color flow image. Newer imaging systems [2] alsodisplay total Doppler power (at the output of the wall filter), P_(d).By not dividing the color flow image by P_(d), we can also display thetrue volume flow, dQ/dt. This is because the numerator,

F=P _(d) ·E(v),

is the power-velocity-integral that is directly proportional to thevolume flow.

Determination of the scale factor, A₀/P₀=A/P₀=dQ/dt/F, that mustmultiply F to obtain volume flow requires further comment

A₀ is the area of a reference beam. In [1], A₀ is smaller than the bloodflow area. We will describe three normalization approaches.

-   -   1. Use a single transmit beam, wider than the vessel, and two        simultaneous receive beams. One receive beam (the measurement        beam) is the same as the transmit beam and the other (the        reference beam) is smaller than the vessel.    -   2. Use two (sequential or multiplexed) two-way (transmit and        receive) beams. One (the measurement beam) is wider than the        vessel and the other (the reference beam is smaller than the        vessel.    -   3. Use two (sequential or multiplexed) two-way (transmit and        receive) beams. Both are wider than the vessel and the        measurement beam is wider than the reference beam.

Let A₀ be the known area of the reference beam, let P₀ and P₁ be themeasured received power in the reference and measurement beams. In case1, the transmit power density is the same for measurement and reference.The receive power is proportional to area. If the area of the vessel (ina slice perpendicular to the line of sight) is A, it follows that

A/A ₀ =P ₁ /P ₀.

In cases 2 and 3, the transmit power density is greater in the referencebeam than in the measurement beam, but by a known factor. In all threecases, the power received in the measurement beam is proportional to thevessel area. In case 2, the received reference power also varies withvessel size, but at a different rate than in the measurement beam. Withproper calibration, correct measurements can be attained in all threecases.

-   [1]. T. Buck, Et al, “Flow Quantification in Valvular Heart Disease    Based on the Integral of Backscattered Acoustic Power Using Doppler    Ultrasound,” Proc. IEEE, vol. 88, no. 3, pp. 307-330, March 2000.-   [2]. K. Ferrara and G DeAngelis, “Color Flow Mapping”, Ultrasound in    Medicine and Biology, vol. 23, no. 2, pp. 321-345, March 1997.

EXAMPLE III 3-D Doppler Ultrasound Blood Flow Monitor with EnhancedField and Sensitivity

This example sets forth an ultrasound Doppler device and method thatenables non-invasive diagnosis (the conventional role of ultrasoundsystems), and also non-invasive unattended and continuous monitoring ofvascular blood flow for medical applications. In particular, theembodiment of the present invention set forth in this example provides:(1) affordable three-dimensional imaging of blood flow using alow-profile easily-attached transducer pad, (2) real-time vectorvelocity, and (3) long-term unattended Doppler-ultrasound monitoring inspite of motion of the patient or pad. None of these three features arepossible with current ultrasound equipment or technology.

The pad and associated processor collects and Doppler processesultrasound blood velocity data in a three-dimensional region through theuse of a two-dimensional phased array of piezoelectric elements on aplanar, cylindrical, or spherical surface. Through use of uniquebeamforming and tracking techniques described herein, the presentinvention locks onto and tracks the points in three-dimensional spacethat produce the locally maximum blood velocity signals. The integratedcoordinates of points acquired by the accurate tracking process is usedto form a three-dimensional map of blood vessels and provide a displaythat can be used to select multiple points of interest for expanded datacollection and for long term continuous and unattended blood flowmonitoring. The three dimensional map allows for the calculation ofvector velocity from measured radial Doppler.

A thinned array (greater than half-wavelength element spacing of thetransducer array) is used to make a device of the present inventioninexpensive and allow the pad to have a low profile (fewer connectingcables for a given spatial resolution) The array is thinned withoutreducing the receiver area by limiting the angular field of view.Grating lobes due to array thinning can be reduced by using widebandwidth and time delay steering. The array, or portions of the array,is used to sequentially insonate the beam positions. Once the region ofinterest has been imaged and coarsely mapped, the array is focused at aparticular location on a particular blood vessel for measurement andtracking. Selection of the point or points to be measured and trackedcan be based on information obtained via mapping and may be user guidedor fully automatic. Selection can be based, for example, on peakresponse within a range of Doppler frequencies at or near an approximatelocation.

In the tracking mode a few receiver beams are formed at a time: sum,azimuth difference, elevation difference, and perhaps, additionaldifference beams, at angles other than azimuth (=0 degrees) andelevation (=90 degrees). Monopulse is applied at angles other than 0 and90 degrees (for example 0, 45, 90, and 135 degrees) in order to locate avessel in a direction perpendicular to the vessel. When the desired(i.e. peak) blood velocity signal is not in the output, this isinstantly recognized (e.g., a monopulse ratio, formed after Dopplerfiltering, becomes non-zero) and the array is used to track (slowmovement) or re-acquire (fast movement) the desired signal.Re-acquisition is achieved by returning to step one to form andDoppler-process a plurality of beams in order to select the beam (andthe time delay or “range gate”) with the most high-Doppler (high bloodvelocity) energy. This is followed by post-Doppler monopulse tracking tolock a beam and range gate onto the exact location of the peak velocitysignal. In applications such as transcranial Doppler, where angularresolution based on wavelength and aperture size is inadequate, finemapping is achieved, for example, by post-Doppler monopulse trackingeach range cell of each vessel, and recording the coordinates andmonopulse-pair angle describing the location and orientation of themonopulse null. With a three-dimensional map available, true vectorvelocity can be computed. For accurate vector flow measurement, themonopulse difference is computed in a direction orthogonal to the vesselby digitally rotating until a line in the azimuth-elevation or C-scandisplay is parallel to the vessel being monitored. The aperture is moreeasily rotated in software (as opposed to physically rotating thetransducer array) if the aperture is approximately circular (oreliptical) rather than square (or rectangular). Also, lower sidelobesresult by removing elements from the four corners of a square orrectangular array in order to make the array an octagon.

All currently available ultrasound devices (including “Doppler colorflow mapping” systems) form images that are limited by their resolution.In some applications, such as TCD, the low frequency required forpenetration of the skull makes the azimuth and elevation resolution atthe depths of interest larger than the vessel diameter. In thisinvention, as long as (1) a blood vessel or (2) a flow region of a givenvelocity can be resolved by finding a 3-D resolution cell through whichonly a single vessel passes, that vessel or flow component can then bevery accurately located within the cell. Monopulse is merely an exampleof one way to attain such sub-resolution accuracy (SRA). Other methodsinvolve “super-resolution” or “parametric” techniques used in “modernspectral estimation”, including the MUSIC algorithm and autoregressivemodeling, for example. SRA allows an extremely accurate map of 3-D flow.

This invention utilizes post-Doppler, sub-resolution tracking andmapping; it does Doppler processing first and uses only highDoppler-frequency data. This results in extended targets since theactive vessels approximate “lines” as opposed to “points”. Inthree-dimensional space, these vessels are resolved, one from another.At a particular range, the monopulse angle axis can be rotated (in theazimuth-elevation plane) so that the “line” becomes a “point” in themonopulse angle direction. That point can then be located by usingsuper-resolution techniques or by using a simple technique such asmonopulse. By making many such measurements an accurate 3-D map of theblood vessels results.

Methods for extending the angular field of view of the thinned array(that is limited by grating lobes) include (1) using multiple panels oftransducers with multiplexed processing channels, (2) convex V-shapedtransducer panels, (3) cylindrical shaped transducer panel, (4)spherical shaped transducer panel, or (5) negative ultrasound lens. Ifneeded, moving the probe and correlating the sub-images can create a mapof an even larger region.

Active digital beamforming can be utilized, but the implementationdepends on a choice to be made between wideband and narrowbandimplementations. If emphasis is on high resolution mapping of the bloodvessels, then a wide bandwidth (e.g., 50% of the nominal frequency) isused for fine range resolution. If emphasis is on Doppler spectralanalysis, measurement, and monitoring, the map is only a tool. In thiscase, a narrowband, low cost, low range-resolution, high sensitivityimplementation might be preferred. A wideband implementation wouldbenefit in performance (higher resolution, wider field of view, andreduced grating lobes) using time-delay steering while a narrowbandimplementation would benefit in cost using phase-shift steering. Theinvention can thus be described in terms of two preferredimplementations.

In a wideband implementation, time delay steering can be implementeddigitally for both transmit and receive by over-sampling and digitallydelaying in discrete sample intervals. In a narrowband implementation,(1) phase steering can be implemented digitally (digital beamforming)for both transmit and receive, and (2) bandpass sampling (sampling at arate lower than the signal frequency) can be employed with digitaldown-conversion and filtering.

Overview of this Embodiment

This embodiment of the present invention involves (1) a family ofultrasound sensors, (2) the interplay of a set of core technologies thatare unique by themselves, and (3) a number of design options whichrepresent different ways to implement the invention. To facilitate anorganizational understanding of this many-faceted invention, adiscussion of each of the three topics above follows.

The sensors addressed are alt two-dimensional (i.e., planar or on thesurface of a convex shape such as a section of a cylinder) arrays ofpiezoelectric crystals for use in active, non-invasive, instantaneous(or real-time), three-dimensional imaging and monitoring of blood flow.While the sensors and the techniques for then use apply to all bloodvessels in the body, the figures and detailed description emphasizes thetranscranial Doppler (TCD) monitor method as a nonlimiting example. Themethod of the present invention utilizes a new, useful and unobviousapproach to 3-D imaging of blood velocity and blood flow that (1) allowsfor finer image resolution than would otherwise be possible with thesame hardware complexity (number of input cables and associatedelectronics) and (2) allows for finer accuracy than would ordinarily bepossible based on the resolution. The invention measures and monitors3-D vector velocity rather than merely the radial component of velocity.

The core technologies that constitute the invention are (1) arraythinning with large elements and limited scanning, (2) array shapes toreduce peak sidelobes and extend the field of coverage. (3) post-Dopplersub-resolution tracking, (4) post-Doppler sub-resolution mapping, (5)additional methods for maximizing the angular field of view, and (6)various digital beamforming procedures for implementing the mapping,tracking, and measurement processes. The invention encompasses arraythinning, where the separation between array elements is significantlylarger than half the wavelength. This reduces the number of input cablesand input signals to be processed while maintaining high resolution andsensitivity and avoiding ambiguities. In the TCD application, wheresignal to noise and hence receiver array area is of paramountimportance, array thinning is possible without reducing the receiverarray area because a relatively small (compared to other applications)angular field of view is needed.

Thinning with full aperture area imposes limitations on the angularfield of view. Methods for expanding the field of view include usingmore elements than are active at any one time. For example, if theelectronics is switched between two identical panels, the cross-rangefield of view at any depth is increased by the size of the panel. If thepanels are pointed in slightly different directions so that overlappingor redundant beams are avoided, the field of view is doubled. Ageneralization of this approach involves the use of an array on acylindrical or spherical surface.

In the TCD application, the achievable angular resolution is poor,regardless of the method of thinning, or whether or not thinning isused. Once a section of a blood vessel is resolved from other vessels inDoppler, depth, and two angles (az and el), Post-Doppler sub-resolutionprocessing locates that section to an accuracy that is one-tenth toone-twentieth of the resolution. This allows for precise tracking andaccurate mapping. Tracking provides for the possibility of unattendedlong term monitoring and mapping aids the operator in selecting thepoint or points to be monitored.

One of ordinary skill in the art will readily recognize that there aremany options available in the design of any member of the family ofsensors that utilizes any or all of the core technologies that comprisethis invention, all of which are encompassed by the present invention, Atwo-dimensional array is established art that can be designed in manyways and can have many sizes and shapes (rectangular, round, etc.).

As with other nonlimiting embodiments of the present invention set forthabove, this embodiment is a non-invasive, continuous, unattended,volumetric, blood vessel tracking, ultrasound monitoring and diagnosticdevice for blood flow. It will enable unattended and continuous bloodvelocity measurement and monitoring as well as 3-dimensional vasculartracking and mapping using an easily attached-electronically steered,transducer probe that can be in the form of a small pad for monitoringapplication, when desired. Although a device of the present inventionhas applications with blood vessels in any part of the body, the cranialapplication will be used as a specific example. A device of the presentinvention can, for example:

-   -   1. Measure and continuously monitor blood velocity with a small        low-profile probe that can be adhered, lightly taped, strapped,        banded, or otherwise easily attached to the portion of the body        where the vascular diagnosis or monitoring is required.    -   2. Track and maintain focus on multiple desired blood vessels in        spite of movement.    -   3. Map 3-D blood flow; e.g., in the Circle of Willis (the        central network of arteries that feeds the brain) or other        critical vessels in the cranial volume.    -   4. Perform color velocity imaging and display a 3-0 image of        blood flow that is rotated via track ball or joystick until a        desired view is selected.    -   5. Form and display a choice of projection, slice, or        perspective views, including (1) a projection on a depth-azimuth        plane, a B-scan, or a downward-looking perspective, (2) a        projection on an azimuth-elevation plane, a C-scan, or a        forward-looking perspective, or (3) a projection on an arbitrary        plane, an arbitrary slice, or an arbitrary perspective.    -   6. Use a track ball and buttons to position circle markers on        the points were measurement or monitoring of vector velocity is        desired.    -   7. Move the track location along the blood vessel by using the        track ball to slide the circle marker along the image of the        vessel.    -   8. Display actual instantaneous and/or average vector velocity,        estimated average volume flow, and/or Doppler spectral        distribution.    -   9. Maintain a multi-day history and display average blood        velocity versus time for each monitored vessel over many hours.    -   10. Sound an alarm when maximum or minimum velocity is exceeded        or when emboli count is high; and maintain a log of emboli        detected.    -   11. Track, map, and monitor small vessels (e.g., 1 mm in        diameter), resolve vessels as close as 4 mm apart (for example),        and locate them with an accuracy of ±0.1 mm, for example.

This embodiment of the present invention will allow a person with littletraining to apply the sensor and position it based on an easilyunderstood ultrasound image display. The unique sensor can continuouslymonitor artery blood velocity and volume flow for early detection ofcritical events. It will have an extremely low profile for easyattachment, and can track selected vessels; e.g., the middle cerebralartery (MCA), with no moving parts. If the sensor is pointed to thegeneral volume location of the desired blood vessel (e.g., within ±1cm.), it will lock to within ±0.1 mm of the point of maximum radialcomponent of blood flow and remain locked in spite of patient movement

A device of the present invention can remain focused on the selectedblood vessels regardless of patient movement because it produces anddigitally analyzes, in real time, a 5-dimensional data base composed ofsignal-return amplitude as a function of:

1. Depth,

2. Azimuth,

3. Elevation,

4. Radial component of blood velocity,

5. Time.

Since a device of the present invention can automatically locate andlock onto the point in three dimensions having the maximum high-Dopplerenergy, i.e. maximum volume of blood having a significant radialvelocity, unattended continuous blood velocity monitoring is one of itsuses. By using the precise relative location of the point at which lockoccurs as a function of depth, a device of the present invention can mapthe network of blood vessels as a 3-dimensional track without thehardware and computational complexity required to form a conventionalultrasound image. Using the radial component of velocity along with thethree-dimensional blood path, a device of the present invention candirectly compute parameters of blood flow, such as vector velocity,blood flow volume, and Doppler spectral distribution.

A device having applications in a method of the present invention is anon-mechanical Doppler ultrasound-imaging sensor comprising probes,processing electronics, and display. Specific choices of probes allowthe system to be used for transcranial Doppler (TCD), cardiac, dialysis,and other applications. Just as with other embodiments of the presentinvention set forth above, this embodiment has application for medicalevaluation and monitoring multiple locations in the body. However, thetranscranial Doppler application will be used as an nonlimiting example.FIG. 13 shows the overall TCD configuration and a typical definition ofthe display screen. The TCD system is comprised of one or two probesthat may be attached to the head with a “telephone operators band” or aVelcro strap. The interface and processing electronics is containedwithin a small sized computer. A thin cable containing from 52 to 120micro coax cables, depending on the example probe design used, attachesthe probe to the electronics in the computer. When the operatorpositions the probe on the head and activates the system, the Anterior,Middle and Posterior Cerebral Arteries and the Circle of Willis aremapped on the screen along with other blood vessels. The arteries orvessels of interest are selected by manually locating a cursor overlaidon the vessel 3-D map. The system locks onto the blood vessels andtracks their position electronically. A variety of selected parametersare displayed on the screen; e.g., the velocity, the pulse rate, depthof region imaged, gain and power level. Using only one probe the TCD canmonitor multiple arteries (vessels) at a time. By way of example,presented on the screen are dual traces, one for each artery selected.The blood velocity can be dynamically monitored. As shown in FIG. 13both the current blood velocity (dark traces) and any historic trace(lighter color) can be displayed simultaneously. The average bloodvelocity or estimated average flow for each artery is displayed belowthe respective velocity trace. The image shows the arteries and thechannel used for each artery. When two probes are used, the display issplit showing signals from both of them. For example, using a differentprobe (i.e., different size) with the same electronics and display, theunit can be used to measure and monitor the blood flow in a carotidartery. Similarly, it can be used to perform this function for dialysis,anesthesia, and in other procedures.

The sensor is a two dimensional array of transducer elements (e.g.,piezoelectric crystals) that are electronically activated in bothtransmit and receive to effect a scan. For example, if a square (N×N)array is used, up to N² elements could be used at the same time. This isillustrated in FIG. 14 for the case of N=8. The array need not besquare. Any M×N array may be utilized in this manner. ALL receivedsignals (52 in the example of FIG. 13) are sampled, digitized, andprocessed. This can be done, for example, in a desk top or lap toppersonal computer with additional cards for electronics and real-timesignal processing as illustrated in FIG. 13 and FIG. 21. The array isphase steered or time-delay steered, depending on the bandwidthutilized, which depends in turn on the desired range resolution. Theangular field of view shown in FIG. 15 is limited by the existence ofgrating lobes caused by array thinning (spacing the array elements morethan ½ wavelength apart). The concept is illustrated below for a1-dimensional array forming a beam that measures only one angle. For atwo-dimensional array, this represents a horizontal or vertical cutthrough the cluster of beams shown in FIG. 15.

The frequency utilized for TCD is usually at or near 2 MHz becausehigher frequencies do not propagate well through bone and lowerfrequencies do not provide adequate reflection from the blood cells.However, other frequencies have applications when examining other padsof the body. With a propagation velocity of 1.54 millimeters permicrosecond, the wavelength is 0.77 mm. If a filled array is utilized,the element size and array pitch would be d=0.77/2. For a cross-rangeresolution of 5.8 mm or less at a depth of 60 mm, the array size, L,must be at least 8 mm (Resolution=depth×wavelength/L). Since N=L/d inFIG. 2, N must exceed 21 and hence the array must have on the order of Nor over 400 elements. If the desired resolution is halved, the arraysize doubles and the number of elements exceeds 1,600. The array in FIG.14 is said to be “thinned” because it only has 52 elements.

As explained above, “grating lobes” are ambiguities or extra, unwanted,beams 35 caused by using a transducer array whose elements are too largeand hence too far apart. The following analysis illustrates grating lobesuppression for the worst case of narrowband signals and phase-shiftbeam processing. Time delay processing using wideband signals would besimilar, but would further attenuate or eliminate grating lobes,resulting in even better performance. Naturally, one of ordinary skillin the art can readily program a processor to suppress or limit gratinglobes with the equations described herein using routine programmingtechniques

Let

x=(d/2)sin θ,  (9)

represent a normalization for the angle, ˜from which reflected acousticenergy arrives. The azimuth (or elevation) angle, θ, is zero in thebroadside direction, perpendicular to the transducer array and d is thewidth (or length) of a single element of the receiver array. Thewavelength of the radiated acoustic wave is λ=c/f where c is theacoustic propagation velocity (1540 meters/second in soft tissue) and fis the acoustic frequency (usually between 1 and 10 megahertz). The widepattern in FIG. 16 a is the element pattern

a _(e)(x)=sin πx/πx.  (10)

The pattern is the product of the element pattern, the array pattern,and cos θ

a(θ)=cos(θ)a _(e)(x)a _(a)(x)  (11)

Each of the two component patterns is plotted separately as a functionof 0 in FIG. 16 a and the total pattern of equation (11) is plotted inFIG. 16 b. In the far-field, i.e., for λr>>L², where r is the range ordepth and L is the length of the aperture, the array pattern steered tothe angle θ=θ₀ is

$\begin{matrix}{{a_{a}(x)} = {\sum\limits_{n = 0}^{N - 1}{w_{n}^{{j2\pi}\; {n{({x - x_{0}})}}}}}} & (12)\end{matrix}$

where w n is a weighting to reduce sidelobes and N is the number ofelements in one dimension. As seen in FIG. 16 a, equation (12) isperiodic in x. The peak at θ=θ₀(θ₀=0 in FIG. 16) is the desired beam andthe others are grating lobes.

In the near field, when focused at (r₀, θ₀) equation (12) is replaced bythe slightly better general Fresnel approximation:

$\begin{matrix}{{a_{a}\left( {x,z} \right)} = {\sum\limits_{n = 0}^{N - 1}{w_{n}^{{j2\pi}{\lbrack{{n{({x - x_{0}})}} + {{({n\; \frac{N - 1}{2}})}^{2}{({z - z_{0}})}}}\rbrack}}}}} & (13)\end{matrix}$

(provided that that the range significantly exceeds the array size,r>L), where x=d sin θ/λ as before, and

z=d ² cos² θ/λr  (14)

Because the receiver aperture is sampled with a spatial period of d, thereceiver array pattern will be periodic in sin θ, with a period of λ/d(equation 12). This periodicity means that the array pattern isambiguous. When the array is pointed broadside (θ=0), it will also bepointed at the angle θ=sin⁻¹ (λ/d), for example. In terms of thenormalized variable, x, the period is unity. Since |sin θ| cannot exceed1, the variable is confined to the interval [−d/λ, d/λ]. Theconventional element spacing is d=λ/2. Thus, in a conventional phasedarray, x is always between −0.5 and +0.5, and hence ambiguities are notencountered. In a highly thinned array (d>λ) there will normally beambiguities or grating lobes as illustrated in FIG. 16 a. The secondgrating lobe, at x=2 or θ=sin⁻¹ (2λ/d) is not real when d does notexceed 2λ

FIG. 16 b shows that the unsteered total pattern does not exhibitgrating lobes. In a 20 two dimensional array, the elements could berectangular instead of square (dx×dy) and the results would still bevalid. Similar results could be obtained for an array in which theelements are staggered from row to row (and/or column to column).

In FIG. 17 the same array is used as in FIG. 16, but the receiverelement signals are combined with a phase taper that steers the beam tox=0.2 or θ=4.71°. In FIG. 17 b, we see that the grating lobes are notcompletely suppressed, with the largest one at x=−1+0.2=−0.8 or0=−19.18°. FIG. 18 shows this in decibels. The worst-case grating lobeis attenuated by at least 12 dB, even in the stressing case of extremelynarrow band operation. These Figures were produced in MATLAB, using thefollowing software (m-file):

MPATTERN xpattern.x Script to plot monostatic patterns vs. theta

Mt=90; wave_length=0.77; d=1.875, N=8,

k=d/wave_length

t=−Mt: 0.1:Mt;

tr=pi.*t·/180;

x=k*sin(tr);

p=pi*x+eps; R=sin(p)·/p;

R=R·*cos(tr);

n=0:N−1;

-   -   xo=0;

xo=0.2; steered

e=exp (j*n′*2*pi*(x−xo));

-   -   w=nanning(N);    -   E=(2/N)*w′*e;

E=(1/N)*ones(1,N)*e;

subplot(211); plot (t, [abs(R);abs (E)]);

ER=abs(E).*abs(R); Monostatic

subplot(212); plot (t, (abs(ER)));

FIG. (2); plot (t, 20*loq 10(abs (ER)));

zoom on;

The values of d and λ used in the above example are representative for atranscranial Doppler application of the invention. If f=2 MHz is chosenfor the center frequency, the wavelength is 0.77 mm. An 8×8 array with awidth and/or length of L=15 mm, provides a one dimensional thinningratio of 2d/λ=4.87. A 15 mm square array with half-wavelength elementswould require more than 15,000 elements. By thinning, this number wasreduced to 52 provided that—the angular field of view is limited to2×4.71=9.42°. For a 1 cm array at 2 MHz, the hyperfocal distance (wherethe 3 dB focal region extends to infinity) is L²/4λ=3.25 cm. For a 15 mmarray, the hyperfocal distance is 7.3 cm. Thus, a fixed focus probesuffices for this application, but the quadratic phase distributionacross the elements required to focus in depth should be added to thelinear phase distributions required to steer the beams.

Using the configuration described above, the cluster of beams in FIG. 15is used to approximately locate the desired point for collecting theblood velocity signal. This is done initially, and is repeatedperiodically, in mapping dwells that are interspersed with normalmeasurement dwells. For example the output of each beam in the clusterwould be Doppler processed by performing an FFT or equivalenttransformation on a sequence of pulse returns. The pulse repetitionfrequency (PRF) would typically be less than or equal to 9 kHz tounambiguously achieve a depth of 8.5 cm for the TCD application. Inorder to obtain a velocity resolution finer than Δv=2 cm per second (todistinguish brain death), a dwell of duration as long as T=λ/(2Δv)=20ms, cr170 pulses at 8.5 kHz, may be desired in the measurement modeDuring monostatic mapping, 21 beams are scanned. If a mapping dwell isto be completed in 20 ms, only 8 pulses per beam are available, and an8-pulse FFT would be utilized for each beam position.

The example shown in FIGS. 16 through 18 was an 8 by 8 receiver arrayforming a 5 by 5 cluster of beams. This is an example of an approximaterule of thumb for this invention, that an N element linear array isrecommended for use in producing N/2+1 beams for N even and [N+1]/2beams for N odd. Thus, a 16 by 10 element rectangular array wouldpreferably be used to form a 9 by 6 cluster of beams, though the actualnumber of beams formed is arbitrary

Because receive beams are formed only in a limited angular region, awide-angle receiver element pattern (which usually implies a smallelement) is not required. In fact, the size of the receiver element canbe as large as the element spacing. Thus the receiver array is “thinned”only in the sense that the element spacing exceeds a half wavelength.Since the element size also exceeds a half wavelength, the array area isnot reduced. It is thinned only in terms of number of elements, not interms of receiver area, Consequently, there is no reduction insignal-to-noise ratio, nor a requirement for increased transmitterpower.

FIG. 19 illustrates a means for increasing the angular field of view inthe azimuth direction by extending the array horizontally. A similarscheme could be used vertically to extend the elevation F.O.V. The52-element array of FIG. 14 becomes a single panel of the extendedarray. One panel is active at a time in FIG. 19. The beamwidth for FIG.14, in radians, is nominally given by λ/L. At a range or depth of R, thecross range resolution is R λ/L (typically 3 to 5 mm). The F.O.V inmillimeters at that same range is less than N/2+1=5 times thatbeamwidth. If a second panel is used in a planar configuration, thesecond panel translates the beam pattern to the right (or left) by thewidth of the panel, L=L₂/2 (typically 8 mm). The field of view can beextended by more than this (it can even be doubled) by tilting the twopanels in opposite directions to minimize the overlap in coverage of thetwo panels.

FIG. 19, with L₁≈L₂, simultaneously provides: (1) a large F.O.V. in theL₂ direction to allow for the simultaneous monitoring of two bloodvessels more than an inch apart, (2) a large active array area for highsensitivity, and (3) a number of active elements below 60 and a totalnumber of elements below 120. An alternative, shown in FIG. 20, has thearray on the surface of a segment of a cylinder. This uses 52 elementsat a time with a total of only 84 elements (and hence only 84 cables).The L₁×L₂′ active array translates around the curved surface as the beamis scanned horizontally. If a symmetric F.O.V. extension (azimuth andelevation) is desired, a spherical surface could be utilized.

FIG. 21 is an overall block diagram depiction of the overall blood flowmonitor. Most functions are performed by means of software in thedigital processor. Naturally, one of ordinary skill in the art canreadily program the processor to perform functions described hereinusing equations set forth herein and routine programming techniques. Apossible implementation of the analog processing is diagrammed in FIG.22. The A/D converter can be a bank of converters or one or moreconverters multiplexed amongst the 52 channels. If an extended arraysuch as shown in FIG. 19 or 20 were used, a switch would be includedbetween the 52 processing channels in FIG. 22 and the actual elements.Note that the 52 element array of FIG. 14 represents an 8×8 array withcorners removed (52=8×8−4×3). Other possibilities include a 24 elementarray (24=6×6−4×3), a 120 element array (120=12×12−4×6), etc.

The transmitter produces pulses for each active element at a pulserepetition frequency (PRF) of 8,500 pulses per second. Each pulse willbe at a frequency of f₀=2 MHz and will have a bandwidth, B, of at least250 kHz (e.g., a pulse no more than 4 microseconds long).

For measurement, only one or two beam positions need be insonated by asingle probe. For mapping, many beam positions must be insonated, withseveral pulses on each for moving target indication (MTI) and/or Dopplerprocessing. A measurement frame duration longer than 20 milliseconds(170 pulses at an 8.5 kHz PRF) may not be necessary because of thenon-stationary (pulsed) nature of human blood flow. Mapping, requiresseveral (4 to 11) pulses per beam position and many (e.g., 21 to 36)beam positions per frame. Since the Doppler resolution for mapping isnot as fine as in the measurement mode, longer mapping frames can beused. If only 21 beams are formed with 8 pulses on each or if up to 34beams are formed with only 5 pulses on each, a frame duration of 20 mscan be maintained ever during search and mapping.

FIG. 22 shows 52 identical receiver chains comprising

-   -   1. Processor controlled time gain control and time gate (open        for up to 26 microseconds for each pulse).    -   2. A limiter for dynamic range control.    -   3. A low noise amplifier (LNA).    -   4. A low pass filter (to pass |f|<f₀+B/2 (e.g., |f|<2.125 kHz)        and reject |f|>5.875 kHz by at least 40 dB (assuming f₀=2 M Hz        and B=250 kHz).

A/D conversion (typically 12 to 16 bits) is performed at an 8 MHz ratefor each channel in FIG. 22. This keeps the analog filteringrequirements extremely simple. It also permits extremely largebandwidths (up to 2 MHz) and time-delay steering. For narrowerbandwidths and phase-shift steering, bandpass analog filtering and muchlower sampling rates (determined by B rather than f₀) could be used. Forthe 8 MHz sampling rate, either time-delay or phase-shift beam steeringcan be utilized (depending on signal bandwidth). FIG. 22 depicts timedelay steering for the transmitter. The distance from each array elementto each focal point (each beam center at a nominal depth (e.g., 60 mmfor TCD) would be pre-computed and stored either as a time delay or as aphase shift (depending on the type of steering) for each element foreach beam. If phase shift steering were utilized on transmit, thetransmitted signal could be created digitally in the processor, followedby D/A conversion for each element. Hence FIG. 22 represents only onepossible embodiment of the invention.

An example of the digital receiver processing for the case of an 8 MHzsampling rate per channel is described below. The input is 208 12 or 16bit samples per pulse (8 samples per microsecond×26 microseconds toallow for a 4 cm deep radial mapping field of view). 8,500pulses/second, and 52 channels. This results in a maximum average rateof 52×208×8500 91.9 MegaSamples per second (or 1.84 million samples in a20 ms frame). During measurement, the range interval can be narrowed toless than 1 cm, reducing the number of samples per pulse to 32. Theaverage rate for measurement and monitoring becomes 14 megasamples persecond. The receiver processing steps are as follows:

Buffer (to allow subsequent processing to be performed at the averagerate).

Digitally Down Convert to Baseband (make I and Q). 52 channels inparallel.

Multiply input samples by samples of a 2 MHz cosine wave and −sine waveto create In-phase and Quadrarture samples, respectively. Since thesamples are ¼ cycle apart, the multiplicands are all 0, 1, or −1, andhence no multiplications are needed. If r(j,p) is the real p^(th) samplefrom the j^(th) channel, the complex low-pass signal, s(j,p) has a realpart for p 0, 1, 2, 3, 4, 5, . . . given by

-   -   r(j,0),0, −r(j,2), 0, r(j, 4),0, . . . .        and an imaginary part given by    -   0, −r(j,1), 0, r(j,3), 0, −r(j,5), . . . .

This provides a data rate 2 times the input rate because the data is nowcomplex.

Pre-Decimation Low-Pass Digital Filter. Filter 52 complex channels. Pass|f|<B/2, reject |f|>r−B/2, where r is the sampling rate after samplerate decimation (e.g., 1 MHz). If B=250 kHz, r could be as low as 500kHz. If B is large, r could be 2 or 3 MHz. If receiver phase-shiftsteering were to be performed, the output samples would be computed atthe decimated rate. If receiver time-delay steering is to be used, weoutput 8 million complex samples per second and postpone sample ratedecimation until after beam formation.

Perform MTI or create coarse Doppler cells. For every channel and everyrange sample, either digitally high-pass filter the sequence of pulsereturns to suppress clutter from tissue and bone or perform 52×2088-point discrete Fourier transforms (DFTs or FFT's) for each mappingframe. (Six points of the 8-point complex DFT provides 3 positive and 3negative coarse Doppler cells.)

Perform Digital Beamforming. Case 1: Time Delay Beamforming with SampleRate Decimation uses a set of pre-computed time delays to reduce 52complex channels with 208 samples per pulse to one of M (e.g. 21)complex beam outputs with 25 samples (range cells) per pulse. Theexample given here assumes 8:1 decimation.

The maximum delay is slightly less than 0.75 μs=6 T, where T=⅛microsecond is the time between input samples. For a given pulse return,the k^(th) sample (k=1, 2, . . . , 25) of the i^(th) beam, i=1, 2, . . ., M, is denoted by b(i, k). The p^(th) sample (p=1, 2, . . . , 208) ofthe j th input channel (=1, 2, . . . , 52) is denoted by s(j,p). Letd_(ij) be the delay required for the signal in channel j to produce beami.

For a given pulse return, the k^(th) complex 1 MHz rate output samplefor beam i is

${b\left( {i,k} \right)} = {\sum\limits_{j = 1}^{52}\left\{ {{a_{ij}{s\left( {j,{{8\left\lbrack {k + 1} \right\rbrack} - b_{ij}}} \right)}} + {\left( {1 - a_{ij}} \right){s\left( {j,{{8\left\lbrack {k + 1} \right\rbrack} - b_{ij} - 1}} \right)}}} \right\}}$

where b_(ij). is the integer part of d_(ij)/T (between 0 and 6) anda_(ij) is the fractional remainder (between 0 and 1). Determine power oramplitude in each output Doppler bin as I²+Q² or its square root:

Case 2: Phase-shift beamforming of already decimated data involves onlya sequence of inner products of 52-dimensional complex vectors ofelement values with a complex vector of representing the required phaseshifts

Display Coarse Blood-Vessel Color-Flow Map. Coarse blood vessel map isthe set of range, azimuth, and elevation cells with high power, with 6Doppler values. Blue and red represent positive and negative Doppler,with saturation related to radial velocity and intensity related topower.

Initialize Acquisition. The user, looking at an azimuth-elevation CoarseMap (with depth automatically truncated to a set of values that shouldinclude the MCA), moves the transducer and looks for a high-intensity,saturated spot. He can center the probe on that spot or he can have adevice of the present invention display a range interval correspondingto the ACA, in which case he can make sure that both vessels are wellwithin the angular field of view of the probe.

Acquisition and Tracking of one or two points being monitored. This isdone with a single transmit beam focused on the spot identified abovefor several frames. Digital Down-conversion, low-pass filtering, and MTIare performed as before, but beamforming is different. Five receivebeams are simultaneously formed. These are a sum beam and four monopulsedifference beams, all steered to the same point as the transmit beam.Each monopulse beam is equivalent to the difference between the outputsof a pair of beams displaced on opposite sides of the focal point. Thefour monopulse pairs are in 45 degree intervals with the first beinghorizontal, and the third being vertical. The monopulse-differenceoutput with the largest magnitude is divided by the output of the sumbeam. The imaginary part is the “monopulse ratio” used to re-steer thebeam (in the difference pair direction) so that it is better centered onthe vessel. This procedure can be repeated in an effort to drive allfour monopulse ratios to zero.

Measurement and Tracking. Tracking continues as described above duringthe measurement mode. Measurement is made with fine Doppler resolution(128 point FFT) applied to only the sum beam. In a 15 ms frame, datafrom 128 pulses are collected (52 channels, 6 range samples). The pulsesare Hamming weighted and FFT'd. This produces 128 Doppler bins (for eachrange bin and element), 66.4 times a second. Real sum beam outputs wouldthen be produced (using monopulse-guided steering) for each of 64 to 126of these Doppler bins.

Ultrasound systems include a beamformer and an image processor. Thetransmit waveforms from the beamformer are converted to acousticalenergy, and the reflected acoustical energy is converted into receivesignals by an array of transducer elements. U.S. application Ser. No.09/926,665, assigned to the present invention and now allowed, describesone method of providing thinned arrays for use in ultrasound systems andis hereby incorporated by reference in its entirety. While the inventionis described in terms of ultrasound applications, and in particular,Doppler ultrasound embodiments, the invention is not limited to DopplerUltrasound operation, but is applicable to all active phased arrays,including sonar, radar, and coherent optics, regardless of whether ornot Doppler processing is involved.

Antifocusing

In imaging blood vessels, it is desirable to obtain a measure of bloodvolume flow in real time. Therefore, the problem of forming N³ voxels inreal time while dwelling long enough to measure blood velocity arises.According to an embodiment of the invention, a solution to this problemis provided by “anti-focusing” the transmitting array. FIG. 24illustrates the concept of antifocusing. For purposes of thisspecification, antifocusing refers to configuring an array 5 to transmita broad transmit beam 6 such that a large region of interest 8, forexample a blood vessel of the human body, is insonated simultaneiously.One method of accomplishing antifocusing the array according to anembodiment of the invention, is by introducing time delays thatcorrespond to a diverging wavefront. Alternative embodiments of theinvention rely on phase shifting to achieve the antifocusing effect. Inany case, one embodiment of the invention employs the concept ofantifocusing a transmitter array. An antifocused transmitter pattern isa pattern that simultaneously illuminates or ensonifies a plurality ofreceiver beams. In one embodiment of the invention, antifocusing isaccomplished by employing time-delays to the transmitter elements. Thetransmitter delays increase with the distance of the array element fromthe center of the array. These delays provide a wavefront that appearsto propagate from the array as if it came from a point source behind thearray. The location of this fictitious point source is the antifocalpoint 9 and its distance from the array center is the antifocaldistance.

Reducing Grating Lobes

A significant disadvantage of known phased arrays is the unwantedpresence of grating lobes and other unpredictable secondary intensitymaxima which can potentially lead to ambiguities in the received signal.The need to reduce side lobes and grating lobes is common to all arraysreported to date. Several techniques including apodization, broadbanding and the use of subsets of elements have been investigated toreduce the effect of side lobes. No effective technique has yet beendeveloped to satisfactorily address the problem of grating lobes. U.S.application Ser. No. 09/926,666 of which this application is acontinuation in part, addresses the problem of grating lobes. Thisphenomenon arises, in part due to the spacing between the elements of atransducer array. It is especially desirable in thinned arrays to reducethe effects of grating lobes. According to an embodiment of theinvention, transmitter patterns are generated that have nulls wherereceiver grating lobes could lead to ambiguities in the received signal.In order to provide such nulls, transmit signal amplitudes are selectedto maintain low transmit sidelobes in the regions of receiver gratinglobes, thereby reducing the effects of the grating lobes.

Simplifying Processing

It is desirable to simplify the processing of received signals and tomaintain the ratio of the received signal to noise and other types ofinterference, including grating lobes. An embodiment of the inventionprovides a transmitter pattern having a nearly constant amplitude over abroader angular region than would be obtained by transmitting a typicallobed pattern. Accordingly, received energy from that region remainsrelatively constant as the illumination of that area remains constantThis simplifies processing of the receive beam information. According toan embodiment of the invention, in order to simplify the processing ofreceived beams, a transmit beam is formed which causes multiple receivearray elements to receive equal reflected energy. FIG. 28 illustratesthe general shape of the transmit beam according to an embodiment of theinvention wherein the nearly constant amplitude portion of the transmitpattern is indicated at 405.

The transmitter pattern illustrated in FIG. 28 is formed by adjustingthe amplitudes of the transmit array elements in addition to theproviding the delays (or phase shifts) that result in antifocuing thebeam. This amplitude adjustment permits the transmit beam to have arelatively flat pattern at the top as well as having nulls in the regionof grating lobes.

The example shown if FIG. 28 is a wide transmit beam patterncharacterized by an almost-constant amplitude almost up to the first dip406, and with a deep null (the second dip 407) occurring at the angle ofthe grating lobes of a center-focused (un-steered) receiver pattern. Oneembodiment of the invention is implemented using a thinned array.

Combination

In one embodiment of the invention, a wide transmit beam pattern ischaracterized by an almost-constant amplitude almost up to the firstdip, and with a deep null (the second dip) occurring at the angle of thegrating lobes of a center-focused (un-steered) receiver pattern using athinned array. By providing a transmit pattern that has been optimizedto provide at substantially all ranges (including the near field), aflat, broad, mainbeam with low sidelobes and a null at the grating lobelocation, multiple receiver elements can be simultaneously ensonated.

FIGS. 25 (a-f) are two dimensional (2-D) Matlab representations of apulsed, digitally processed, sampled and quantized transmitter beampattern according to an embodiment of the invention. In FIG. 25 range isdenoted by the letter “r”. The x axis of each figure indicates crossrange in millimeters. The y axis in each figure indicates transmitteroutput amplitude in decibels (db). The location of a would be receivergrating lobe corresponding to the cross range direction of thetransmitted energy is indicated by an arrow in each figure. The patternis illustrated at six different ranges (also referred to herein asdepths): 15 mm, 20 mm, 30 mm, 40 mm, 60 mm and 120 mm respectively. Theexample shown represents a demodulated signal for a pulse comprising aplurality of cycles of a 6 Mhz carrier. As can be seen from the drawing,the waveform is characterized by having nulls (indicated by arrows) inpositions corresponding to receiver grating lobes.

FIGS. 26 (a-f) are Matlab pulsed simulations of patterns of a 16 elementlinear receive array. The simulations plot amplitude as a function ofacoustic scatter cross range location, according to an embodiment of theinvention. The figures compare the center focused array patterns (a-c)to the patterns (d-f) of an array steered 3 elements to the left ofcenter. In FIGS. 26 (a-c) the pattern is center focused at x₁, z₁=(0 mm,60 mm) with a resolution of 1.2 mm. FIG. 26 a illustrates the amplitudeof the combined transmit/receive pattern in the x, z plane. FIG. 26 b isa 2-D linear representation of the combined transmit/receive signalillustrated in FIG. 2 a plotting amplitude as a function of cross rangeposition in mm. In FIG. 26 b cross range is indicated in mm along thex-axis. And amplitude is indicated along the y-axis. In FIGS. 26 a, b, dand e amplitude is plotted linearly, e.g., volts. In figures FIGS. 26 cand f, amplitude is plotted logarithmically in db. FIGS. 26 b and cillustrate sidelobes at 211 and 220 and main lobe at 210. FIG. 26 c is a2-D logarithmic representation of the receiver and transmitter patternsillustrated in FIG. 26 b. FIG. 26 c illustrates sidelobes 220 and 225 of−33 db on either side of main lobe 210.

FIG. 26 d illustrates the amplitude of the combined transmit/receivepattern in the x, z plane of the same array shown in 26 a, moved threeelements to the left, i.e., (x₁, z₁) (−2.7 mm, 60 mm). FIGS. 26 e and26F show in 2-D the corresponding linear and logarithmic representationsof the receiver gain pattern as a function of acoustic scatter crossrange location at resolution 1.3 mm. As shown in FIG. 26 f, a −25 dbgrating lobe 230 appears at approximate cross range 22 mm. Grating lobe230 can also be seen in FIG. 26 e at 230.

FIGS. 27 (a-f) represent a simplified 2-D continuous wave (CW)simulation of a transmitter pattern according to an embodiment of theinvention. Transmitter amplitude (in db) at ranges of 20 mm, 30 mm, 40mm, 50 mm, 60 mm and 70 mm respectively are illustrated. The pattern ischaracterized by a main lobe 310 having a generally flat top shape. Thisshape allows multiple receive beams to be received at the same relativeamplitudes, thereby simplifying processing of the received information.The shape also allows the ratio of Signal to (Noise+Interference) to bemaintained independently of receive beam direction.

FIG. 28 illustrates the 3-D pattern 400 of a 2-D array according to theembodiment of the invention illustrated in FIGS. 27 (a-d). Thetransmitter array is anti-focused and amplitude weighted as describedherein to produce a pattern 400 characterized by having a main lobe 410with a generally flat top portion 405 such that an N×N array of receivebeams will be illuminated with equal amplitude. Pattern 400 is furthercharacterized by having nulls 407 in the region where grating lobes ofthe receive beams would normally appear.

The beam shapes illustrated in FIG. 27 is described by the relationshiph(u)=c for |u|<u₁ and h(u)=0 for |u|>u₂, where u=sin θ. In this manner,several receive elements can be ensoni fled within the transmitterpattern while grating lobes are attenuated. At short range (in the nearfield), the transmited beam shape is approximately the same as thereceived illumination function. In this case very few transmit elementsare entirely on. At long range (far field) the receive elementillumination function g(x) or g(y) and the transmitted pattern, h(u) area Fourier transform pair. Thus, one embodiment of the invention uses asinc function (sin x/x) for illumination. The desired illuminationfunction is found by performing a Fast Fourier Transform (an FFT) on asequence consisting of several constant positive values centered aboutzero (with zeros elsewhere), and then ignoring every second term of theFFT. For example, using the centered odd FFT outputs as weightingfactors to be applied to array elements provides the four-element lineararray weights,

-   -   [−0.2, 1.0, 1.0, −0.2]

Wherein the array weight numbers represent amplitudes of the energyapplied to the transmit elements.

This example weighting factor set produces desired patterns according tothe invention at long ranges and at short ranges. In one embodiment ofthe invention, in addition to the weighting factors given above, delaysare applied to the inner elements corresponding to focusing the array atabout 20 mm (delay=[0, 3, 3, 0] in tics of a 96 MHz clock). As will berecognized by those of ordinary skill in the art, other combinations ofweights and delays can also be used to produce the pattern of theinvention. One embodiment of the invention, employing time delaysteering, provides a 180-degree phase shift for the outside elements byeither delaying or advancing the outer elements by half a cycle. Thiscorresponds, respectively, to either anti-focusing the array, orfocusing it at an extremely short range. In this embodiment, the delaythat corresponds to a 180 degree phase shift for a 6 MHz carrier, is 8ticks of the 96 MHz clock. Accordingly, the element weights become

-   -   [0.2, 1.0, 1.0, 0.2]        and the antifocusing delays become [5, 0, 0, 5] (this is        obtained by subtracting 3 from [8, 3, 3, 8] (measured in tics of        the 96 MHz clock) ([8, 3, 3, 8] is obtained by adding 8 to the        outside elements of the [0,3,3,0] delay described above. In one        embodiment of the invention good overall performance, using        non-negative weights, was obtained with the weights above and        delays of [6, 0, 0, 6] at 96 MHz. These delays correspond to        focusing about 9 mm behind the array (in other words,        anti-focusing).

In the [5, 0, 0, 5] embodiment anti-focuses of −10 mm instead of −9 mmalso produced the waveform of the invention. By adding 8 to [−8, 3, 3,−8] delays of [0, 11, 11, 0] are obtained. This corresponds to focusingat a very short range. At ranges of interest, the result is the same asantifocusing

FIG. 29 is a waveform diagram representing one of many implementationsof the waveforms used to create the transmitter beam profiles of theinvention. Various embodiments of the invention will use more or lesselements to create the transmitted beam. The example shown employs an8×8 array. The numbers appearing from left to right across the top ofFIG. 29 represent increments of time units. The waveforms (0-7) enableenergy transmission from a corresponding element when high and disableenergy transmission from a corresponding element when low. In theexample shown time starts with 0, so waveform 0 has a delay of zero.

Waveform 2 has a delay of 25 and so forth. Transmitter elements usingwaveform 1 are always disabled. Each waveform 1-7 transmits acorresponding number of pulses. For example, waveform 0 transmits 4pulses, waveform 2 transmits 1 cycle and so on. The relative widths ofthe pulses high indicate the duty cycle of the transmitter. For example,waveform 1 represents a 50% duty cycle, and waveform 5 represents a12.5% duty cycle. The initial positive going edge of each waveformrelative to time zero indicates the delay. For example, waveform 0 has adelay of 0 time units and waveform has a delay of 8 time units. Thevariation in duty cycle and number of pulses are used in lieu ofamplitude variation.

Table 1 illustrates the waveform assignments for an 8×8 array ofelements according to one embodiment of the invention.

$\begin{matrix}\begin{matrix}1 & 1 & 1 & 1 & 1 & 1 & 1 & 1 \\1 & 1 & 7 & 5 & 5 & 7 & 1 & 1 \\1 & 7 & 4 & 3 & 3 & 4 & 7 & 1 \\1 & 5 & 3 & 0 & 0 & 3 & 5 & 1 \\1 & 5 & 3 & 0 & 0 & 3 & 5 & 1 \\1 & 7 & 4 & 3 & 3 & 4 & 7 & 1 \\1 & 1 & 7 & 5 & 5 & 7 & 1 & 1 \\1 & 1 & 1 & 1 & 1 & 1 & 1 & 1\end{matrix} & {{TABLE}\mspace{20mu} 1}\end{matrix}$

Table 2 illustrates the contents of a Read only memory storing thewaveform

Sel ROM cycles High Low Delay Start 0 waveform 0 4 8  8 0 0X3B 1waveform 1 — — — — 2 waveform 2 1 2 14 25  0X54 3 waveform 3 4 4 12 40X3F 4 waveform 4 4 3 13 6 0X41 5 waveform 5 4 2 14 8 0X43 6 waveform 62 2 14 25  0X54 7 waveform 7 2 2 14 26  0X55

Various publications are cited herein, the disclosures of which areincorporated by reference in their entireties.

1. A method for long term Doppler ultrasound monitoring using an arrayof ultrasound transducer elements, comprising: collecting and Dopplerprocessing ultrasound blood velocity data in a three dimensional region;locking onto and tracking the point(s) in three-dimensional space thatproduce the locally maximum blood velocity signals.
 2. A method of claim1, additionally comprising measuring radial velocity at the point(s). 3.A method of claim 1, additionally comprising forming a three-dimensionalmap of at least one blood vessel based on integrating coordinates ofpoints acquired by the tracking process.
 4. A method of claim 1, whereinthe ultrasound blood velocity data is radial Doppler data, andadditionally comprising calculating vector velocity from the Dopplerultrasound blood velocity data.
 5. A method of claim 1, additionallycomprising providing a display that allows selection of multiple pointsof interest for expanded data collection.
 6. A method of claim 1,additionally comprising forming a broad ultrasound transmit beamencompassing a plurality of narrow receive beams and initially acquiringthe blood velocity data by insonating a large region.
 7. A method ofclaim 1, additionally comprising applying post Doppler sub-resolutionprocessing to locate the point(s) being tracked to an accuracy that isfiner than the resolution.
 8. A method of claim 1, additionallycomprising tracking and maintaining focus on multiple blood vessels. 9.A method of claim 1, additionally comprising performing color velocityimaging.
 10. A method of claim 1, additionally comprising forming anddisplaying a choice of projection, slice or perspective views.
 11. Amethod of claim 1, additionally comprising displaying at least one ofinstantaneous blood flow velocity and average blood flow velocity andestimated average blood flow volume.
 12. A method of claim 1,additionally comprising maintaining a multi-day history and displayingaverage blood flow velocity versus time for each monitored vessel over atime period.
 13. A method of claim 1, additionally comprising soundingan alarm when at least one of a maximum blood flow velocity, a minimumblood flow velocity, and an emboli count is high.
 14. A method of claim1, additionally comprising electronically rotating the array or portionsof the array using digital processing techniques.
 15. A transcranialDoppler device for long term Doppler ultrasound monitoring comprising anarray of ultrasound transducer elements and processing electronics thatproduce and digitally analyze data composed of ultrasound signal returnamplitude as a function of depth, azimuth, elevation, radial bloodvelocity and time and that automatically locates and locks onto thepoint(s) with the maximum volume of blood having a significant radialvelocity.
 16. An ultrasound device comprising an array of ultrasoundtransducer elements and additionally comprising processing electronicsprogrammed to operate the array differently in transmit and receivemodes, to form a transmit beam encompassing a plurality of receive beamsfor initially acquiring a signal by insonating a target regioncomprising multiple receive beam positions simultaneously, to receiveand Doppler process data from the multiple receive beam positions of thearray, and to lock onto and track the point(s) in three-dimensionalspace that produce the locally maximum blood velocity signals.
 17. Adevice of claim 16, wherein the processing electronics are programmed tosteer the receive beams using a phase steering or time-delay steeringtechnique.
 18. A device of claim 16, wherein the processing electronicsare additionally programmed to correct for motions in the target regionby periodically forming multiple receive beams and re-acquiring a peaksignal.
 19. A device of claim 16, wherein the array of ultrasoundtransducer elements is provided on a low-profile easily-attachedtransducer pad.
 20. A device of claim 16, wherein the processingelectronics are additionally programmed to determine spatial coordinatesof received data.
 21. A device of claim 20, additionally comprising adisplay device, and wherein the processing electronics are additionallyprogrammed to form and display a 3D map based on the spatial coordinatesof received data.